Implantable pulse generator systems and methods for operating the same

ABSTRACT

Improved assemblies, systems, and methods provide a stimulation system for prosthetic or therapeutic stimulation of muscles, nerves, or central nervous system tissue, or any combination. The stimulation system includes an implantable pulse generator and a lead sized and configured to be implanted subcutaneously in a tissue region. An external controller includes circuitry adapted for wireless telemetry and a charging coil for generating the radio frequency magnetic field to transcutaneously recharge a rechargeable battery in the pulse generator. Using wireless telemetry, the pulse generator is adapted to transmit status information back to the external controller to allow the external controller to automatically adjust up or down the magnitude of the radio frequency magnetic field and/or to instruct a user to reposition the charging coil, the status information adapted to allow optimal recharging of the pulse generator rechargeable battery.

RELATED APPLICATIONS

This application is a continuation of U.S. application Ser. No.11/517,213, filed Sep. 7, 2006, granted as U.S. Pat. No. 9,308,382,which claims the benefit of U.S. Provisional Patent Application Ser. No.60/801,003, filed 17 May 2006, and entitled “Implantable Pulse Generatorfor Providing Functional and/or Therapeutic Stimulation of Muscle and/orNerves and/or Central Nervous System Tissue.” The entire content of eachof these applications is incorporated herein by reference.

U.S. application Ser. No. 11/517,213 is a continuation-in-part of U.S.patent application Ser. No. 11/149,654, filed 10 Jun. 2005, granted asU.S. Pat. No. 7,565,198, and entitled “Systems and Methods for BilateralStimulation of Left and Right Branches of the Dorsal Genital Nerves toTreat Dysfunctions, Such as Urinary Incontinence,” which claims thebenefit of U.S. Provisional Patent Application Ser. No. 60/578,742,filed Jun. 10, 2004, and entitled “Systems and Methods for BilateralStimulation of Left and Right Branches of the Dorsal Genital Nerves toTreat Dysfunctions, Such as Urinary Incontinence.” The entire content ofeach of these applications is incorporated herein by reference.

U.S. application Ser. No. 11/517,213 is also a continuation-in-part ofU.S. patent application Ser. No. 11/150,418, filed 10 Jun. 2005, grantedas U.S. Pat. No. 7,239,918, and entitled “Implantable Pulse Generatorfor Providing Functional and/or Therapeutic Stimulation of Musclesand/or Nerves and/or Central Nervous System Tissue,” which claims thebenefit of U.S. Provisional Patent Application Ser. No. 60/599,193,filed Aug. 5, 2004, and entitled “Implantable Pulse Generator forProviding Functional and/or Therapeutic Stimulation of Muscles and/orNerves,.” The entire content of each of these applications isincorporated herein by reference.

U.S. application Ser. No. 11/517,213 is also a continuation-in-part ofU.S. patent application Ser. No. 11/150,535, filed 10 Jun. 2005, grantedas U.S. Pat. No. 7,813,809, and entitled “Implantable Pulse Generatorfor Providing Functional and/or Therapeutic Stimulation of Musclesand/or Nerves and/or Central Nervous System Tissue,” which claims thebenefit of U.S. Provisional Patent Application Ser. No. 60/680,598,filed May 13, 2005, and entitled “Implantable Pulse Generator forProviding Functional and/or Therapeutic Stimulation of Muscles and/orNerves and/or Central Nervous System Tissue.” The entire content of eachof these applications is incorporated herein by reference.

FIELD OF THE INVENTION

The invention relates to systems and methods for providing stimulationof central nervous system tissue, muscles, or nerves, or combinationsthereof.

BACKGROUND OF THE INVENTION

Neuromuscular stimulation (the electrical excitation of nerves and/ormuscle to directly elicit the contraction of muscles) andneuromodulation stimulation (the electrical excitation of nerves, oftenafferent nerves, to indirectly affect the stability or performance of aphysiological system) and brain stimulation (the stimulation of cerebralor other central nervous system tissue) can provide functional and/ortherapeutic outcomes. While existing systems and methods can provideremarkable benefits to individuals requiring neuromuscular orneuromodulation stimulation, many limitations and issues still remain.For example, existing systems often can perform only a single, dedicatedstimulation function.

Today there are a wide variety of implantable medical devices that canbe used to provide beneficial results in diverse therapeutic andfunctional restorations indications. For example, implantable pulsegenerators can provide therapeutic and functional restoration outcomesin the field of urology, such as for the treatment of (i) urinary andfecal incontinence; (ii) micturition/retention; (iii) restoration ofsexual function; (iv) defecation/constipation; (v) pelvic floor muscleactivity; and/or (vi) pelvic pain. Implantable pulse generators can alsobe used for deep brain stimulation, compensation for various cardiacdysfunctions, pain management by interfering with or blocking painsignals, vagal nerve stimulation for control of epilepsy, depression, orother mood/psychiatric disorders, the treatment of obstructive sleepapnea, for gastric stimulation to prevent reflux or to reduce appetiteor food consumption, and can be used in functional restorationsindications such as the restoration of motor control.

There exists both external and implantable devices for providingbeneficial results in diverse therapeutic and functional restorationsindications. The operation of these devices typically includes the useof an electrode placed either on the external surface of the skin, avaginal or anal electrode, or a surgically implanted electrode. Althoughthese modalities have shown the ability to provide a neurologicalstimulation with positive effects, they have received limited acceptanceby patients because of their limitations of portability, limitations oftreatment regimes, and limitations of ease of use and user control.

Implantable devices have provided an improvement in the portability ofneurological stimulation devices, but there remains the need forcontinued improvement. Implantable stimulators described in the art haveadditional limitations in that they are challenging to surgicallyimplant because they are relatively large, they require direct skincontact for programming and for turning on and off, and only provide asingle dedicated stimulation function. In addition, current implantablestimulators are expensive, owing in part to their limited scope ofusage.

These implantable devices are also limited in their ability to providesufficient power which limits their use in a wide range of stimulationapplications, requires surgical replacement of the device when thebatteries fail, and limits their acceptance by patients. Rechargeablebatteries have been used but are limited by the need to recharge a powersupply frequently (e.g., daily), and the inconvenience of specialrecharge methods.

More recently, small, implantable microstimulators have been introducedthat can be injected into soft tissues through a cannula or needle.Although these small implantable stimulation devices have a reducedphysical size, their application to a wide range of simulationapplications is limited. Their micro size extremely limits their abilityto maintain adequate stimulation strength for an extended period withoutthe need for frequent recharging of their internal power supply(battery). Additionally, their very small size limits the tissue volumesthrough which stimulus currents can flow at a charge density adequate toelicit neural excitation. This, in turn, limits or excludes manyapplications.

For each of these examples, the medical device (i.e., an implantablepulse generator), is often controlled using microprocessors withresident operating system software (code). This operating systemsoftware may be further broken down into subgroups including systemsoftware and application software. The system software controls theoperation of the medical device while the application software interactswith the system software to instruct the system software on what actionsto take to control the medical device based upon the actual applicationof the medical device (i.e., to control incontinence or the restorationof a specific motor control).

As the diverse therapeutic and functional uses of implantable medicaldevices increases, and become more complex, system software having aversatile interface is needed to play an increasingly important role.This interface allows the system software to remain generally consistentbased upon the particular medical device, and allows the applicationsoftware to vary greatly depending upon the particular application. Aslong as the application software is written so it can interact with theinterface, and in turn the system software, the particular medicaldevice can be used in a wide variety of applications with only changesto application specific software. This allows a platform device to bemanufactured in large, more cost effective quantities, with applicationspecific customization occurring at a later time.

It is time that systems and methods for providing neurologicalstimulation address not only specific prosthetic or therapeuticobjections, but also address the quality of life of the individualrequiring the beneficial stimulation. In addition, there remains theneed for improved size, operation, and power considerations ofimplantable medical devices that will improve the quality of life issuesfor the user.

SUMMARY OF THE INVENTION

The invention provides improved assemblies, systems, and methods forproviding prosthetic or therapeutic stimulation of central nervoussystem tissue, muscles, or nerves, or muscles and nerves.

One aspect of the invention provides an implantable pulse generatorsystem. The system comprises a pulse generator including a housing sizedand configured for implantation in subcutaneous tissue, circuitrycarried within the housing and adapted for wireless telemetry, arechargeable battery coupled to the circuitry and carried within thehousing, the battery having a voltage, a power receiving coil carriedwithin the housing and coupled to the circuitry, the power receivingcoil for transcutaneously receiving an externally generated radiofrequency magnetic field to recharge the rechargeable battery, and anexternal controller including circuitry adapted for wireless telemetry,and a charging coil for generating the radio frequency magnetic field totranscutaneously recharge the rechargeable battery.

Another aspect of the invention provides a pulse generator system wherethe power receiving coil includes a maximum outside dimension X, andwherein the pulse generator is adapted to be implanted in subcutaneoustissue at an implant depth D, such that the ratio of X/D is betweenabout 0.8 to 1 and about 4 to 1. The outer surface of the housingmaintains a two degrees Celsius or less temperature rise during the timeperiod in which the power receiving coil is transcutaneously receivingthe externally generated radio frequency magnetic field. In oneembodiment, the implant depth D is between about five millimeters andabout twenty millimeters. The housing of the implantable pulse generatormay be sized to have a thickness of between about 5 mm and 15 mm, awidth of between about 30 mm and 60 mm, and a length of between about 45mm and 60 mm, and the circuitry carried within the housing is operablefor generating electrical stimulation pulses and/or sensing myoelectricsignals.

During a battery recharge period, and using wireless telemetry, thepulse generator is adapted to transmit status information back to theexternal controller to allow the external controller to automaticallyadjust up or down the magnitude of the radio frequency magnetic fieldand/or to instruct a user to reposition the charging coil, the statusinformation adapted to allow optimal recharging of the pulse generatorrechargeable battery. The instruction to the user may be a visualinstruction and/or an audio instruction. The adjustment of the magnitudeof the radio frequency magnetic field may be an adjustment of up toabout 300 percent of the initial magnitude. The status information caninclude an indication of the battery charge status and an indication ofthe magnitude of power recovered by the receive coil. The statusinformation can also include a magnitude of the battery voltage and amagnitude of a DC voltage recovered from the radio frequency magneticfield. The radio frequency magnetic field comprises a frequency betweenabout 30 KHz and about 300 KHz.

Another aspect of the invention provides an implantable pulse generatorsystem where the pulse generator is exclusively the wireless telemetrycommunications slave, with all wireless telemetry communicationsinitiated by the external controller. The wireless telemetrycommunication between the external controller and the pulse generatoroperates in the MICS (Medical Implant Communications Service) band atbetween about 402 MHz to about 405 MHz.

Yet another aspect of the invention provides an implantable pulsegenerator system where the external controller's charging coil may becoupled via a cable to the external controller. The charging coilcomprises an outside diameter in the range of about 40 millimeters toabout 70 millimeters, and a thickness as measured perpendicular to thediameter in the range of about three millimeters to about elevenmillimeters. The charging coil is adapted to maintain a temperature ator below about 41 degrees Celsius during a charging period.

Yet another aspect of the invention includes an implantable pulsegenerator system where the rechargeable battery of the implantable pulsegenerator comprises a capacity of at least 30 mA-hr and recharging ofthe rechargeable battery is required less than weekly. When therechargeable battery has only a safety margin charge remaining, it canbe recharged in a time period of not more than six hours.

Another aspect of the invention provides an implantable pulse generatorsystem, the system comprising a housing sized and configured forimplantation in subcutaneous tissue, circuitry carried within thehousing and adapted for wireless telemetry, a rechargeable batterycoupled to the circuitry and carried within the housing, a powerreceiving coil carried within the housing and coupled to the circuitry,the power receiving coil for transcutaneously receiving an externallygenerated radio frequency magnetic field to recharge the rechargeablebattery, and an external controller including circuitry adapted forwireless telemetry, and a charging coil for generating the radiofrequency magnetic field to transcutaneously recharge the rechargeablebattery.

The pulse generator wireless telemetry circuitry includes a transceiverchip, the transceiver chip adapted to be off about 99 percent or more ofthe time for efficient power management, and is pulsed on periodicallyto search for a command from the external controller. The wirelesstelemetry is configured to have the external controller issue commandsto the pulse generator at timed intervals to confirm that the generatedradio frequency magnetic field is adequate for recharging therechargeable battery.

The circuitry within the housing further includes power managementcircuitry, the power management circuitry comprising at least threeoperating modes relating to the operation of the pulse generator, theoperating modes including pulse generator active, pulse generatordormant, and pulse generator active and charging. The circuitry alsoincludes a microcontroller, the microcontroller adapted to generate aserial data stream, the serial data stream being converted by the pulsegenerator wireless telemetry circuitry into a pulsing carrier signalduring a transmit process, and the pulse generator wireless telemetrycircuitry adapted to convert a varying radio frequency signal strengthinto a serial data stream during a receive process.

Still another aspect of the invention provides an implantable pulsegenerator system, the system comprising a pulse generator including ahousing sized and configured for implantation in subcutaneous tissue,circuitry carried within the housing and adapted for wireless telemetry,a rechargeable battery coupled to the circuitry and carried within thehousing, and a power receiving coil carried within the housing andcoupled to the circuitry, the power receiving coil for transcutaneouslyreceiving an externally generated radio frequency magnetic field torecharge the rechargeable battery. The system also comprises an externalcontroller including circuitry, a rechargeable battery coupled to thecircuitry, and the circuitry adapted for wireless telemetry, and acharging coil coupled to the external controller for generating theradio frequency magnetic field to transcutaneously recharge therechargeable battery. During a rechargeable battery recharge period, theexternal controller is adapted to be carried by a user with noconnection to a power main to allow the user to be completely mobile.

Another aspect of the invention provides an implantable pulse generatorincluding a stimulation component having an on state to generate astimulation pulse and an off state, a receiver component selectivelyenabled to receive information using wireless telemetry, and acontroller coupled to the stimulation component and the receivercomponent to enable operation of the receiver component only when thestimulation component is in the off state.

Yet another aspect of the invention provides an implantable pulsegenerator including a case, a rechargeable battery positioned within thecase to provide operating power, charge management circuitry positionedwithin the case and coupled to the rechargeable battery, the chargemanagement circuitry implementing a constant current phase and aconstant voltage phase charge regime, a controller positioned within thecase and coupled to the charge management circuitry, the controller fortiming the constant voltage phase of the rechargeable battery chargeregime, and a receive coil positioned in the case and coupled to therechargeable battery, the receive coil receiving externally generatedenergy from a recharger to recharge the rechargeable battery.

The pulse generator is adapted to recharge a fully dischargedrechargeable battery in less than about six hours. The controllerfurther monitors the rechargeable battery voltage and terminates thecharge regime after the rechargeable battery voltage has been in theconstant voltage phase for greater than a predetermined time period. Thecontroller terminates the charge regime by wireless telemetrycommunication with the recharger and instructing the recharger to stoprecharging.

Another aspect of the invention provides a method of wireless telemetrycommunication of status information between an implanted pulse generatorand an external controller including the steps of providing a pulsegenerator, the pulse generator comprising a housing sized and configuredfor implantation in subcutaneous tissue, circuitry carried within thehousing and adapted for wireless telemetry, a rechargeable batterycoupled to the circuitry and carried within the housing, the batteryhaving a voltage, and a power receiving coil carried within the housingand coupled to the circuitry, the power receiving coil fortranscutaneously receiving an externally generated radio frequencymagnetic field to recharge the rechargeable battery.

The steps further including providing an external controller includingcircuitry adapted for wireless telemetry, and a charging coil forgenerating the radio frequency magnetic field to transcutaneouslyrecharge the rechargeable battery, and during a battery recharge periodand using wireless telemetry, the pulse generator transmitting statusinformation back to the external controller to allow the externalcontroller to automatically adjust up or down the magnitude of the radiofrequency magnetic field and/or to instruct a user to reposition thecharging coil, the status information adapted to allow optimalrecharging of the pulse generator rechargeable battery. The instructionto the user may be a visual instruction and/or an audio instruction.

The steps may further may further include providing an instruction sheetto the user, the instruction sheet defining how to reposition thecharging coil for optimal recharging based on the visual and/or audioindication provided by the external controller.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a diagrammatic view of a stimulation system that provideselectrical stimulation to central nervous system tissue, muscles and/ornerves inside the body using a general purpose implantable pulsegenerator, the system including internal and external components thatembody the features of the invention.

FIG. 2A is an anatomical view showing an implantable pulse generatorwith a lead and electrode implanted in tissue.

FIG. 2B is a side view showing a representative implant depth of theimplantable pulse generator in tissue.

FIGS. 3A and 3B are front and side views of the general purposeimplantable pulse generator as shown in FIG. 1, which is powered by arechargeable battery.

FIGS. 3C and 3D are front and side views of an alternative embodiment ofa general purpose implantable pulse generator as shown in FIG. 1, whichis powered using a primary battery.

FIG. 4A is a perspective view of the general purpose implantable pulsegenerator as shown in FIG. 1, without a lead and electrode.

FIG. 4B is an exploded view of the implantable pulse generator as shownin FIG. 4A, showing the general components that make up the implantablepulse generator.

FIG. 4C is a section view of the receive coil taken generally along line4C-4C in FIG. 4B.

FIG. 4D is a top plan view of the receive coil shown in FIG. 4C, showingthe maximum outside dimension.

FIGS. 5 through 15 are perspective views showing possible steps forassembling the implantable pulse generator shown in FIG. 4B.

FIG. 16 is a perspective view of the smaller end of the implantablepulse generator shown in FIG. 4A prior to assembling the header to theimplantable pulse generator.

FIG. 17 is a perspective view of the implantable pulse generator duringa vacuum bake-out process and prior to assembling the header.

FIG. 18 is a perspective view of the implantable pulse generator duringthe backfill and welding process and prior to assembling the header.

FIG. 19 is a perspective view of the implantable pulse generator shownin FIG. 4A with the header positioned for attachment.

FIG. 20 is a diagrammatic view showing operating system software beingdownloaded to the implantable pulse generator using wireless telemetry.

FIG. 21 is a perspective view of the implantable pulse generator shownin FIG. 4A, including a lead and electrode.

FIG. 22A is an anatomic view showing the implantable pulse generatorshown in FIGS. 3A and 3B having a rechargeable battery and shown inassociation with a transcutaneous implant charger controller (batterycharger) including a separate, cable coupled charging coil whichgenerates the RF magnetic field, and also showing the implant chargercontroller using wireless telemetry to communicate with the implantablepulse generator during the charging process.

FIG. 22B is an anatomic view showing the transcutaneous implant chargercontroller as shown in FIG. 22A, including an integral charging coilwhich generates the RF magnetic field, and also showing the implantcharger controller using wireless telemetry to communicate with theimplantable pulse generator.

FIG. 22C is a perspective view of the implant charger controller of thetype shown in FIGS. 22A and 22B, with the charger shown connected to thepower mains to recharge the power supply within the implant chargercontroller.

FIG. 23A is an anatomic view showing the implantable pulse generatorshown in FIGS. 3A through 3D in association with a clinical programmerthat relies upon wireless telemetry, and showing the programmer'scapability of communicating with the implantable pulse generator up toan arm's length away from the implantable pulse generator.

FIG. 23B is a system view of an implantable pulse generator systemincorporating a network interface and showing the system's capability ofcommunicating and transferring data over a network, including a remotenetwork.

FIG. 23C is a graphical view of one possible type of patient controllerthat may be used with the implantable pulse generator shown in FIGS. 3Athrough 3D.

FIG. 24 is a block diagram of a circuit that the implantable pulsegenerator shown in FIGS. 3A and 3B may utilize.

FIG. 25 is an alternative embodiment of the block diagram shown in FIG.24, and shows a block circuit diagram that an implantable pulsegenerator shown in FIGS. 3C and 3D and having a primary battery mayutilize.

FIG. 26A is a circuit diagram showing a possible circuit for thewireless telemetry feature used with the implantable pulse generatorshown in FIGS. 3A through 3D.

FIG. 26B is a graphical view of the wireless telemetry transmit andreceive process incorporated in the circuit diagram of FIG. 26A.

FIG. 27 is a circuit diagram showing a possible circuit for the stimulusoutput stage and output multiplexing features used with the implantablepulse generator shown in FIGS. 3A through 3D.

FIG. 28 is a graphical view of a desirable biphasic stimulus pulseoutput of the implantable pulse generator for use with the system shownin FIG. 1.

FIG. 29 is a circuit diagram showing a possible circuit for themicrocontroller used with the implantable pulse generator shown in FIGS.3A through 3D.

FIG. 30 is a circuit diagram showing one possible option for a powermanagement sub-circuit where the sub-circuit includes MOSFET isolationbetween the battery and charger circuit, the power managementsub-circuit being a part of the implantable pulse generator circuitshown in FIG. 24.

FIG. 31 is a circuit diagram showing a second possible option for apower management sub-circuit where the sub-circuit does not includeMOSFET isolation between the battery and charger circuit, the powermanagement sub-circuit being a part of the implantable pulse generatorcircuit shown in FIG. 24.

FIG. 32 is a circuit diagram showing a possible circuit for the VHHpower supply feature used with the implantable pulse generator shown inFIGS. 3A through 3D.

FIG. 33 is a perspective view of the lead and electrode associated withthe system shown in FIGS. 1 and 2A.

FIGS. 34A and 34B are side interior views of representative embodimentsof a lead of the type shown in FIG. 33.

FIG. 35 is an end section view of the lead taken generally along line35-35 in FIG. 34A.

FIG. 36 is an elevation view, in partial section, of a lead andelectrode of the type shown in FIG. 33 residing within an introducersheath for implantation in a targeted tissue region, the anchoringmembers being shown retracted within the sheath.

FIG. 37 is a perspective view of a molded cuff electrode positionedabout a target nerve N.

FIG. 38 is a diagrammatic view of the custom operating system software,including system software and application software.

FIG. 39 is an anatomic view showing the long lead length feature of theimplantable pulse generator, the lead capable of extending an anatomicalfurthest distance to deliver electrical stimulation.

The invention may be embodied in several forms without departing fromits spirit or essential characteristics. The scope of the invention isdefined in the appended claims, rather than in the specific descriptionpreceding them. All embodiments that fall within the meaning and rangeof equivalency of the claims are therefore intended to be embraced bythe claims.

DESCRIPTION OF THE PREFERRED EMBODIMENT

The various aspects of the invention will be described in connectionwith providing stimulation of central nervous system tissue, muscles, ornerves, or muscles and nerves for prosthetic or therapeutic purposes.That is because the features and advantages that arise due to theinvention are well suited to this purpose. Still, it should beappreciated that the various aspects of the invention can be applied toachieve other objectives as well.

I. The Implantable Pulse Generator System

FIG. 1 shows in diagrammatic form an implantable pulse generator system10. The implantable pulse generator system 10 can be used forstimulating a central nervous system tissue, nerve, or a muscle, or anerve and a muscle to achieve a variety of therapeutic (treatment) orfunctional (restoration) purposes.

The implantable pulse generator system 10 may include both implantablecomponents and external components. The implantable components mayinclude, but are not limited to: an implantable pulse generator 18coupled to a lead 12 and an electrode 16. The external components mayinclude, but are not limited to: a clinical programmer 108, aprint/backup station 110, a docking station 107, a network interface 116(external controller derivative), an implant charger controller 102, acharging coil 104, a power adapter 106, a patient controller 114, aninstruction sheet 120, and a magnet 118. Each of these components of thesystem 10 will be described in greater detail below.

As an exemplary embodiment, the implantable pulse generator may be usedto provide therapeutic restoration for urinary urge incontinence bystimulation of afferent nerves. In this application, a sequence (regime)of nerve stimulation is provided to maintain a level of nervous systemmediation that prevents spasms of the bladder-sensory reflex. Thepredefined stimulus regime may include: a programmable period of nostimulation (a gap), a transition from no stimulation to fullstimulation (ramp up), a period of constant, full stimulation (burst),and transition back to no stimulation (ramp down). This cycle repeatsindefinitely; except as may be modified by a clinician or patientrequest for higher or lower stimulus strength. That request may be madeusing a clinical programmer 108, the implant charger controller 102, orthe patient controller 114, for example, using the wireless telemetry112. Instructions 120 may be provided to describe operation and usagefor all components and all users (i.e., clinician and patient).

A. Implantable Pulse Generator Components

FIG. 2A shows the implantable pulse generator 18 coupled to theimplantable lead 12. The distal end of the lead 12 includes at least oneelectrically conductive surface, which will in shorthand be called anelectrode 16. The electrode 16 may also be positioned along the lengthof the lead 12. The electrode 16 is implanted in electrical conductivecontact with at least one functional grouping of nerve tissue, muscle,or at least one nerve, or at least one muscle and nerve, depending onthe desired functional and/or therapeutic outcome desired. The lead 12,electrode 16, and the implantable pulse generator 18 are shown implantedwithin a tissue region T of a human or animal body.

The implantable pulse generator 18 is housed within an electricallyconductive titanium case or housing 20 which can also serve as a returnelectrode for the electrical stimulus current introduced by thelead/electrode when operated in a monopolar configuration. Theimplantable pulse generator 18 includes a connection header 26 thatdesirably carries a plug-in receptacle for the lead 12. In this way, thelead 12 electrically connects the electrode 16 to the implantable pulsegenerator 18. The case 20 is desirably shaped with a smaller end 22 anda wider end 24, with the header 26 coupled to the smaller end 22. AsFIG. 2A shows, this geometry allows the smaller end 22 of the case 20(including the header 26), to be placed into the skin pocket P first,with the wider end 24 being pushed in last.

The implantable pulse generator 18 is sized and configured to beimplanted subcutaneously in tissue, desirably in a subcutaneous pocketP, which can be remote from the electrode 16, as FIG. 2A shows. Theimplantable pulse generator 18 is capable of driving large electricalresistance occurring in long lead lengths, e.g., the lead 12 is capableof extending an anatomical furthest distance. The anatomical furthestdistance may be the full length of the body; from head to toe in ahuman. For example, the implantable pulse generator could be implantedin an upper chest region and the lead could extend down to the foot (seeFIG. 39). This capability allows the implantable pulse generatorplacement to be selected conveniently and not be constrained by thelocation of the electrode.

In order to accomplish driving the generated electrical stimulationcurrent or pulses from the implantable pulse generator 18 through thelead 12 extending the anatomic furthest distance, the implantable pulsegenerator includes a software programmable VHH power supply 134 (to bedescribed in greater detail later) that can produce the necessary highervoltages. This power supply is software programmable to provide avoltage large enough to drive the requested stimulation current throughthe lead 12 and electrode 16 circuit resistance/impedance. The VHH powersupply 134 can be adjusted up to about 27 VDC. This relatively largevoltage allows the delivery of cathodic phase currents up to about 20 mAinto long lead lengths or into higher impedance electrodes.

In an exemplary application, (an intramuscular stimulating electrode 16with the case 20 as the return electrode, for example), the total tissueaccess resistance of the electrode-to-tissue interface is between about100 ohms and 500 ohms. The lead 12 connecting the electrode(s) 16 to theimplantable pulse generator 18 have resistances that are roughlyproportional to the length of the lead. Typical leads have resistancesin the range of about 2 ohms to 5 ohms of electrical resistance forevery centimeter of lead length. Thus, a relatively long lead, 70 cm forexample, may have about 350 ohms of lead resistance. Combined with about500 ohms of tissue access resistance, this gives a total patient circuitresistance of up to about 850 ohms. To drive 20 mA through this circuit,the VHH power supply 134 would be programmed to provide about 17 VDC.

Desirably, the implantable pulse generator 18 is sized and configured tobe implanted using a minimally invasive surgical procedure. The surgicalprocedure may be completed in a number of steps. For example, once alocal anesthesia is established, the electrode 16 is positioned at thetarget site. Next, a subcutaneous pocket P is made and sized to acceptthe implantable pulse generator 18. A finger dissection, e.g., theclinician's thumb, for example, may be used to form the pocket P afteran initial incision has been made. The pocket P is formed remote fromthe electrode 16. Having developed the subcutaneous pocket P for theimplantable pulse generator 18, a subcutaneous tunnel is formed forconnecting the lead 12 and electrode 16 to the implantable pulsegenerator 18. The lead 12 is routed through the subcutaneous tunnel tothe pocket site P where the implantable pulse generator 18 is to beimplanted. The lead 12 is then coupled to the implantable pulsegenerator 18, and both the lead 12 and implantable pulse generator 18are placed into the subcutaneous pocket, which is sutured closed.

FIG. 4B shows an exploded view of the implantable pulse generator 18shown in FIG. 4A. As shown in FIG. 4B, the case 20 includes a bottomcase component 28 and a top case component 30. Within the bottom case 28and top case 30 is positioned a circuit 32 for generating the electricalstimulation waveforms. An on-board, primary or rechargeable battery 34desirably provides the power. The implantable pulse generator 18 alsodesirably includes an on-board, programmable microcontroller 36, whichcarries operating system code. The code expresses pre-programmed rulesor algorithms under which the desired electrical stimulation waveformsare generated by the circuit 32.

According to its programmed rules, when switched on, the implantablepulse generator 18 generates prescribed stimulation waveforms throughthe lead 12 and to the electrode 16. These stimulation waveformsstimulate the central nervous system tissue, muscle, nerve, or bothnerve and muscle tissue that lay in electrical conductive contact (i.e.,within close proximity to the electrode surface where the currentdensities are high) with the electrode 16, in a manner that achieves thedesired therapeutic (treatment) or functional restoration result.Examples of desirable therapeutic (treatment) or functional restorationindications will be described in greater detail in section III.

Within the case 20 is also positioned a bottom nest 38 and a top nest40. The plastic nests 38 and 40 provide support for the circuitry 32, aweld band 37, and a receive coil 42. A number of feed-thrus 44, 46, 48are coupled to the bottom case 28 and/or top case 30 and provideelectrical connectivity between the circuitry within the case and aheader 26 while maintaining the hermetic seal of the case. The header 26is positioned over the feed-thrus 44, 46, 48 at the smaller end 22 ofthe case 20.

1. Implantable Pulse Generator Assembly

A representative process for assembling the implantable pulse generator18 will now be described. It is to be appreciated that the process forassembling the implantable pulse generator 18 is not intended to belimiting, but merely an example to describe the interrelation of theimplantable pulse generator 18 components shown in FIG. 4B. As FIGS. 5and 6 shows, the feed-thrus 44, 46, 48 are coupled (e.g., welded orbraised), to preexisting apertures in the bottom case 28 and top case30. As shown, feed-thru 44 and 46 are coupled to the bottom case 28 andfeed-thru 48 is coupled to the top case 30.

As shown in FIG. 4B, feed-thru 48 is coupled to the wireless telemetryantenna 80. The antenna 80 may be a conductor separate from conductor 60(see FIG. 7), or it may be the same conductor. If a separate conductoris used (for example because a metal with better electrical conductivityis deemed desirable for operation of the antenna), then there will be acoupling between the two conductors (60 & 80). It is likely that thiscoupling will be a crimp connection or a weld, although not limited toonly these coupling configurations.

Each feed-thru 44, 46, 48, includes a feed-thru conductor 64, 62, 60respectively, to be coupled to the circuitry 32 and the header 26. FIG.7 shows feed-thru 48 in detail. As can be seen, a conductor 60 passesthrough a glass or ceramic insulator 66 of the feed-thru.

The circuitry 32 is sized and configured to precisely fit within the topnest 40 and bottom nest 38, which in turn precisely fit within the topcase 30 and bottom case 28. As can be seen in FIG. 8A, the circuitry 32first comprises a generally flat configuration using flexible circuitboard technology. The circuitry 32 comprises a top circuit portion 50electrically coupled to a bottom circuit portion 52 by way of a flexiblehinge portion 53. The top circuit 50 includes an antenna tab 54 and alead tab 56. The bottom circuit 52 includes a battery tab 58. In orderto fit the circuitry 32 within the case 20, the bottom circuit 52 isfolded over the top circuit 50 and the battery tab 58 is folded inwardtoward the top circuit 50, as can be seen in FIG. 8B. The battery 34 maythen be positioned and coupled (e.g., soldered), to the inward facingbattery tab 58. The lead tab 56 may then be folded upward and inwardtoward the bottom circuit 52, and the antenna tab 54 may be foldedupward and inward toward the bottom circuit, as can be seen in FIG. 8C.The circuitry 32, including the battery 34, may now be positioned withinthe bottom case 28 and top case 30.

The bottom case 28 and the top case 30 may be positioned in a fixture(not shown) to aid with the assembly process. The antenna tab 54 and thelead tab 56 are electrically coupled to their respective feed-thrus inthe bottom case 28 and top case 30.

As shown in FIG. 9, conductor 60 of feed-thru 48 is coupled to theantenna tab 54, and conductors 62 and 64 of feed-thrus 46 and 44respectively are coupled to lead tab 56. Lead tab 56 is also coupled toa ground pin 59 coupled to the inside of the bottom cover 28.

Next, the top nest 40 is positioned within the top case 30 (see FIG.10). The circuitry 32 is then positioned within the top case 30 and thetop nest 40. The receive coil 42 is then seated within the top nest 40and electrically coupled to the circuitry 32 (see FIGS. 11 and 12). Thebottom nest 38 is then seated over the receive coil 42 and the circuitry32 (see FIG. 13), and the weld band 37 is secured over the top nest 40and bottom nest 38 (see FIG. 14). A “getter” 35 may be positioned withinthe bottom case 28 and the top case 30 at any time prior to putting thecase pieces together. The getter 35 helps to eliminate any moisture orother undesirable vapors that may remain in the case 20 after the casehas been sealed. The bottom case 28 can then be positioned on the topcase 30 (see FIGS. 15 and 16).

Next, the assembled implantable pulse generator 18 is subjected to avacuum bake-out process in chamber 70 (see FIG. 17). The vacuum bake-outprocess drives out any moisture content within the unsealed implantablepulse generator 18 and drives out any other volatile contaminants inpreparation for the final sealing of the implantable pulse generator 18.After a predetermined bake-out period (e.g., 45 degrees Celsius to 100degrees Celsius, and for 24 to 48 hours), the chamber 70 is thenbackfilled with an inert gas or gas mixture 72, such as helium-argon(see FIG. 18). A laser welder 74 then applies a weld 76 to the seam 78where the bottom case 28 and top case 30 come together. The weld band 37protects the components within the case 20 during the laser weldingprocess.

A final assembly process may include coupling the header 26 to thesmaller end 22 of the case 20 and the exposed electrical conductors 60,62, 64 (see FIGS. 16 and 19). The header 26 includes connector blocksfor the IS-1 connector inserted or molded within. The header 26 also hasslots or passages molded within for holding the antenna 80, an antennainsert 81, the conductors 62 and 64 of feed-thrus 44 and 46, and theheader brackets 98 and 99 (see FIG. 4B). The thin plastic antenna insert81 is used to guide the bending of the antenna 80 and to secure theantenna 80 inside the header 26.

With the antenna 80 bent around the antenna insert 81, and the otherfeed-thru conductors 62 and 64 sticking out straight, the header 26 isslipped onto the flat face of the welded case (the flat face as shown inFIG. 16). The antenna 80, the antenna insert 81, the feed-thruconductors 62 and 64, and the header brackets 98 and 99, all slip intoslots or passages molded into the header 26 as the header fits flushagainst the case. The feed-thru conductors 62 and 64 are then welded tothe connector blocks inside the header through slots or apertures moldedin the header. Anchor pins 94 and 96 are slipped through the apertures98 and 99 in the header brackets 90 and 92 and into anchor pin slots orapertures molded into the header 26. The anchor pins 94 and 96 arewelded to the header brackets 90 and 92 and mechanically secure theheader to the case through the header brackets.

Any remaining space between the header 26 and the case 20 may also bebackfilled with an adhesive, such as silicone, to seal the header 26 tothe case 20 and fill any remaining gaps. Similarly, the holes throughwhich the anchor pins were installed and the holes through which thefeed-thru conductors were welded to the connector blocks are alsobackfilled with adhesive, such as silicone. The final result is ahermetically sealed implantable pulse generator 18, as seen in FIGS. 20and 21.

FIG. 20 also shows programming the implantable pulse generator 18 withoperating system software, system software, and/or application software.A programmer 84 may be used to download system software, which may ormay not include the application software, to the implantable pulsegenerator 18. This feature of programming, or reprogramming, theimplantable pulse generator 18 allows the implantable pulse generator tobe manufactured and partially or fully programmed. The implantable pulsegenerator may then be put into storage until it is to be implanted, oruntil it is known what application software is to be installed. Thedownloading of the application software or changes to the applicationsoftware can take place anytime prior to implantation. This featuremakes use of a set of software which was programmed into themicrocontroller during the manufacturing process. The programmer 84 maybe similar to the clinical programmer 108 or a modified clinicalprogrammer, except with added features to allow for the programming orreprogramming of the implantable pulse generator 18.

B. Implantable Pulse Generator Features

Desirably, the size and configuration of the implantable pulse generator18 makes possible its use as a general purpose or universal device(i.e., creating a platform technology), which can be used for manyspecific clinical indications requiring the application of pulse trainsto central nervous system tissue, muscle and/or nervous tissue fortherapeutic (treatment) or functional restoration purposes. Most of thecomponents of the implantable pulse generator 18 are desirably sized andconfigured so that they can accommodate several different indications,without major change or modification. Examples of components thatdesirably remains unchanged for different indications include the case20, the battery 34, the power management circuitry 130, themicrocontroller 36, much of the operating system software (firmware) ofthe embedded code, and the stimulus power supply (VHH and VCC). Thus, anew indication may require only changes to the programming of themicrocontroller 36. Most desirably, the particular code may be remotelyembedded in the microcontroller 36 after final assembly, packaging, andsterilization of the implantable pulse generator 18.

Certain components of the implantable pulse generator 18 may be expectedto change as the indication changes; for example, due to differences inleads and electrodes, the connection header 26 and associatedreceptacle(s) for the lead may be configured differently for differentindications. Other aspects of the circuit 32 may also be modified toaccommodate a different indication; for example, the stimulator outputstage(s), or the inclusion of sensor(s) and/or sensor interfacecircuitry for sensing myoelectric signals.

In this way, the implantable pulse generator 18 is well suited for usefor diverse indications. The implantable pulse generator 18 therebyaccommodates coupling to a lead 12 and an electrode 16 implanted indiverse tissue regions, which are targeted depending upon thetherapeutic (treatment) or functional restoration results desired. Theimplantable pulse generator 18 also accommodates coupling to a lead 12and an electrode 16 having diverse forms and configurations, againdepending upon the therapeutic or functional effects desired. For thisreason, the implantable pulse generator can be considered to be generalpurpose or “universal.”

1. Desirable Technical Features

The implantable pulse generator 18 can incorporate various technicalfeatures to enhance its universality.

a. Small, Composite Case

According to one desirable technical feature, the implantable pulsegenerator 18 can be sized small enough to be implanted (or replaced)with only local anesthesia. As FIGS. 3A and 3B show, the functionalelements of the implantable pulse generator 18 (e.g., circuit 32, themicrocontroller 36, the battery 34, and the connection header 26) areintegrated into a small, composite case 20. As can be seen, the case 20defines a small cross section; e.g., about (5 mm to 12 mm thick)×(15 mmto 40 mm wide)×(40 mm to 60 mm long). The overall weight of theimplantable pulse generator 18 may be approximately eight to fifteengrams. These dimensions make possible implantation of the case 20 with asmall incision; i.e., suitable for minimally invasive implantation.Additionally, a larger, and possibly similarly shaped implantable pulsegenerator might be required for applications with more stimulus channels(thus requiring a large connection header) and or a larger internalbattery.

FIGS. 3C and 3D illustrate an alternative embodiment 88 of theimplantable pulse generator 18. The implantable pulse generator 88utilizes a primary battery 34. The implantable pulse generator 18 sharesmany features of the primary cell implantable pulse generator 88. Likestructural elements are therefore assigned like numerals. As can be seenin FIGS. 3C and 3D, the implantable pulse generator 88 may comprise acase 20 having a small cross section, e.g., about (5 mm to 15 mmthick)×(45 mm to 60 mm wide)×(45 mm to 60 mm long). The overall weightof the implantable pulse generator 88 may be approximately fifteen tothirty grams These dimensions make possible implantation of the case 20with a small incision; i.e., suitable for minimally invasiveimplantation.

The case 20 of the implantable pulse generator 18 is desirably shapedwith a smaller end 22 and a larger end 24. As FIG. 2A shows, thisgeometry allows the smaller end 22 of the case 20 to be placed into theskin pocket P first, with the larger end 22 being pushed in last.

As previously described, the case 20 for the implantable pulse generator18 comprises a laser welded titanium material. This construction offershigh reliability with a low manufacturing cost. The clam shellconstruction has two stamped or successively drawn titanium case halves28, 30 that are laser welded around the internal components andfeed-thrus 44, 46, 48. The molded plastic spacing nests 38, 40 is usedto hold the battery 34, the circuit 32, and the power recovery (receive)coil 42 together and secure them within the titanium case 20.

As can be seen in FIG. 2B, the implantable pulse generator 18 may beimplanted at a target implant depth of not less than about fivemillimeters beneath the skin, and not more than about twenty millimetersbeneath the skin, although this implant depth may change due to theparticular application, or the implant depth may change over time basedon physical conditions of the patient. The targeted implant depth is thedepth from the external tissue surface to the closest facing surface ofthe implantable pulse generator 18.

The thickness of the titanium for the case 20 is selected to provideadequate mechanical strength while balancing the greater powerabsorption and shielding effects to the low to medium frequency magneticfield 100 used to transcutaneously recharge the implantable rechargeablebattery 34 with thicker case material (the competing factors are poortransformer action at low frequencies—due to the very low transferimpedances at low frequencies—and the high shielding losses at highfrequencies). The selection of the titanium alloy and its thicknessensures that the titanium case allows adequate power coupling torecharge the secondary power source (described below) of the implantablepulse generator 18 at the target implant depth using a low to mediumfrequency radio frequency (RF) magnetic field 100 from an implantcharger controller 102 and associated charging coil 104 positioned overor near the implantable pulse generator 18 (see FIGS. 22A and 22B).

b. Internal Power Source

According to one desirable technical feature, the implantable pulsegenerator 18 desirably possesses an internal battery capacity or chargesufficient to allow operation with a recharging duty cycle of not morefrequently than once per week for many or most clinical applications.The battery 34 of the implantable pulse generator 18 desirably can berecharged in less than approximately six hours with a rechargingmechanism that allows the patient to sleep in bed or carry on mostnormal daily activities while recharging the battery 34 of theimplantable pulse generator 18. The implantable pulse generator 18desirably has a service life of greater than three years with thestimulus being a high duty cycle, e.g., virtually continuous, lowfrequency, low current stimulus pulses, or alternatively, the stimulusbeing higher frequency and amplitude stimulus pulses that are used onlyintermittently, e.g., a very low duty cycle.

To achieve this feature, the battery 34 of the implantable pulsegenerator 18 desirably comprises a secondary (rechargeable) powersource; most desirably a Lithium Ion battery 34. Given the averagequiescent operating current (estimated at 8 μA plus 35 μA for a wirelesstelemetry receiver pulsing on twice every second) and a seventy percentefficiency of the stimulus power supply, a 1.0 Amp-hr primary cellbattery can provide a service life of less than two years, which is tooshort to be clinically or commercially viable for most indications.Therefore, the implantable pulse generator 18 desirably incorporates asecondary battery, e.g., a Lithium Ion rechargeable battery that can berecharged transcutaneously. Given representative desirable stimulationparameters (which will be described later), a Lithium Ion secondarybattery with a capacity of at least 30 mA-hr will operate for over threeyears. Lithium Ion implant grade batteries are available from a domesticsupplier. A representative battery capacity for one embodiment having acapacity of up to four stimulus channels provides about 130 to about 250milliWatt-hr (approximately 30 milliAmp-hr to 65 milliAmp-hr) in apackage configuration that is of appropriate size and shape to fitwithin the implantable pulse generator 18. For an alternative embodimenthaving a capacity of eight or more stimulus channels, a representativebattery capacity provides about 250 to about 500 milliWatt-hr(approximately 66 milliAmp-hr to 131 milliAmp-hr).

The implantable pulse generator 18 desirably incorporates circuitryand/or programming to assure that the implantable pulse generator 18will suspend stimulation at a first remaining battery capacity and asthe remaining capacity decreases, eventually suspend all operations whenonly a safety margin of battery capacity remains. For example, theimplantable pulse generator 18 may be adapted to suspend stimulation atthe first remaining battery capacity (e.g., about fifteen percent toabout thirty percent of battery capacity remaining), and perhapsfall-back to only very low rate telemetry, and eventually suspends alloperations when the battery 34 has reached the safety margin, i.e., asecond remaining battery capacity (e.g., about five percent to abouttwenty percent of battery capacity remaining). At this second remainingbattery capacity, the battery 34 has discharged the majority of itscapacity, described as a fully discharged battery, and only the safetymargin charge remains. Once in this Dormant mode, the implantable pulsegenerator 18 is temporarily inoperable and inert. The safety margincharge ensures that the implantable pulse generator may be able toremain in the Dormant mode and go without recharging for at least sixmonths. A delay in recharging for at least six months will not causepermanent damage or permanent loss of capacity to the lithium battery34. If the battery 34 goes without charging for much longer than sixmonths, the battery's self-discharge may cause a loss of batterycapacity and/or permanent damage.

The power for recharging the battery 34 of the implantable pulsegenerator 18 is provided through the application of a low frequency(e.g., 30 KHz to 300 KHz) RF magnetic field 100 applied by a skin orclothing mounted implant charger controller 102 placed over or near theimplantable pulse generator (see FIGS. 22A and 22B). The implant chargercontroller 102 might use a separate RF magnetic coupling coil (chargingcoil) 104 which is placed and/or secured on the skin or clothing overthe implantable pulse generator 18 and connected by cable to the implantcharger controller 102 (circuitry and battery in a housing) that is wornon a belt or clipped to the clothing (see FIG. 22A). In an alternativeapplication, it is anticipated that the user would wear the implantcharger controller 102, including an internal RF magnetic coupling coil(charging coil) 104, over the implantable pulse generator 18 to rechargethe implantable pulse generator 18 (see FIG. 22B). The implant chargercontroller 102 allows the patient the freedom to move about and continuewith most normal daily activities while recharging the implantable pulsegenerator.

The charging coil 104 preferably includes a predetermined construction,e.g., desirably 150 to 250 turns, and more desirably 200 turns of sixstrands of #36 enameled magnetic wire (all six strands being wound nextto each other and electrically connected in parallel), or the like.Additionally, the charging coil outside diameter is in a range of about40 millimeters to about 70 millimeters, and desirably about 65millimeters, although the diameter may vary. The thickness of thecharging coil 104 as measured perpendicular to the mounting plane is tobe significantly less than the diameter, e.g., about three millimetersto about eleven millimeters, so as to allow the coil to be embedded orlaminated in a sheet to facilitate placement on or near the skin. Such aconstruction will allow for efficient power transfer and will allow thecharging coil 104 to maintain a temperature at or below about 41 degreesCelsius.

The implant charger controller 102 preferably includes its own internalbatteries which may be recharged from the power mains, for example. Apower adapter 106 may be included to provide for convenient rechargingof the system's operative components, including the implant chargercontroller and the implant charger controller's internal batteries (seeFIG. 22C). The implant charger controller 102 may not be used torecharge the implantable pulse generator 18 while plugged into the powermains.

Desirably, the implantable pulse generator 18 may be recharged while itis operating and the outer surface of the case 20 will not increase intemperature by more than two degrees Celsius above the surroundingtissue during the recharging. It is desirable that for most applicationsthe recharging of the fully discharged battery 34 requires not more thansix hours, and a recharging would be required between once per month toonce per week depending upon the power requirements of the stimulusregime used.

c. Wireless Telemetry

According to one desirable technical feature, the assembly or system 10includes an implantable pulse generator 18, which desirably incorporateswireless telemetry (rather that an inductively coupled telemetry) for avariety of functions able to be performed within arm's reach of thepatient, the functions including receipt of programming and clinical(e.g., stimulus) parameters and settings from the clinical programmer108, communicating usage history and battery status to the clinicalprogrammer, providing user control of the implantable pulse generator18, and for controlling the RF magnetic field 100 generated by theimplant charger controller 102.

Each implantable pulse generator may also have a unique signature,(e.g., a serial number, which may include a model and/or series number,stored in non-volatile memory), that limits communication (securecommunications) to only the dedicated controllers (e.g., the matchedimplant charger controller 102, patient controller 114, or a clinicalprogrammer 108 configured with the serial number for the implantablepulse generator in question). The clinical programmer may be configuredfor use (i.e., wireless telemetry) with many patients by configuring theclinical programmer with a desired serial number to select a specificimplantable pulse generator.

The implantable pulse generator 18 desirably incorporates wirelesstelemetry as an element of the implantable pulse generator circuit 32shown in FIG. 24. A circuit diagram showing a desired configuration forthe wireless telemetry feature is shown in FIG. 26A. It is to beappreciated that modifications to this circuit diagram configurationwhich produce the same or similar functions as described are within thescope of the invention.

As shown in FIG. 23A, the system 10 desirably includes an externalcontroller, such as the clinical programmer 108 that, through a wirelesstelemetry 112, transfers commands, data, and programs into theimplantable pulse generator 18 and retrieves status and data out of theimplantable pulse generator 18. In some configurations, the clinicalprogrammer may communicate with more than one implantable pulsegenerator implanted in the same user. Timing constraints imposed on theexternal controller and the implantable pulse generator 18 prevents twoor more implantable pulse generators or two or more external controllersfrom communicating at nearly the same time. This eliminates thepossibility that a response from one implantable pulse generator will bemisinterpreted as the response from another implantable pulse generator.

The clinical programmer 108 initiates the wireless telemetrycommunication 112(1) to the implantable pulse generator 18, thecommunication including the implantable pulse generator's unique serialnumber and data elements that indicate the communication is a commandfrom an external controller, e.g., data elements in a packet header.Only the implantable pulse generator 18 having the unique serial numberresponds 112(2) to the clinical programmer's communication. Thecommunication response 112(2) includes data elements that indicate thecommunication is a response to a command from an external controller,and not a command from an external controller.

An external controller such as the clinical programmer 108 may alsoinclude provisions to seek out implantable pulse generators withincommunication range without knowing a unique serial number. Toaccomplish this, the clinical programmer may search for a range ofserial numbers, such as 1 to 1000, as a non-limiting example.

The clinical programmer 108 may incorporate a custom programmed generalpurpose digital device; e.g., a custom program, industry standardhandheld computing platform or other personal digital assistant (PDA).The clinical programmer 108 can also include an on-board microcontrollerpowered by a rechargeable battery. The rechargeable battery of theclinical programmer 108 may be recharged when connected via a cable tothe print/backup station 110, or docked on the docking station 107 (acombined print/backup station and recharge cradle)(see FIG. 1). Inaddition to recharging the battery of the clinical programmer, thedocking station 107 and/or the print/backup station 110 may also providebackup, retrieve, and print features. The docking station 107 and/or theprint/backup station 110 may include memory space to allow the clinicalprogrammer to download or upload (via wireless communication, a cable,and/or a portable memory device) any and all information stored on theclinical programmer 108 (backup and retrieve feature), and also allowthe information from the clinical programmer 108 to be printed in adesired format (print feature).

In addition, the rechargeable battery of the clinical programmer 108 maybe recharged in the same or similar manner as described and shown inFIG. 22C for the implant charger controller 102, i.e., connected to thepower mains with a power adapter 106 (see FIG. 1); or the customelectronics of the clinical programmer 108 may receive power from theconnected pocket PC or PDA.

The microcontroller carries embedded code which may includepre-programmed rules or algorithms that allow a clinician to remotelydownload program stimulus parameters and stimulus sequences parametersinto the implantable pulse generator 18. The microcontroller of theclinical programmer 108 is also desirably able to interrogate theimplantable pulse generator and upload usage data from the implantablepulse generator. FIG. 23A shows one possible application where theclinician is using the programmer 108 to interrogate the implantablepulse generator. FIG. 23B shows an alternative application where theclinical programmer, or a network interface 116 intended for remoteprogramming applications and having the same or similar functionality asthe clinical programmer 108 or the implant charger controller 102, isused to interrogate the implantable pulse generator. As can be seen, thenetwork interface 116 is connected to a local computer, allowing forremote interrogation via a local area network, wide area network, orInternet connection, for example.

Using subsets of the clinical programmer software, features of theclinical programmer 108 or network interface 116 may also include theability for the clinician or physician to remotely monitor and adjustparameters using the Internet or other known or future developednetworking schemes. The network interface 116 would desirably connect tothe patient's computer in their home through an industry standardnetwork such as the Universal Serial Bus (USB), where in turn an appletdownloaded from the clinician's server would contain the necessary codeto establish a reliable transport level connection between theimplantable pulse generator 18 and the clinician's client software,using the network interface 116 as a bridge. Such a connection may alsobe established through separately installed software. The clinician orphysician could then view relevant diagnostic information, such as thehealth of the unit or its current settings, and then modify the stimulussettings in the implantable pulse generator or direct the patient totake the appropriate action. Such a feature would save the clinician,the patient and the health care system substantial time and money byreducing the number of office visits during the life of the implant.

Other features of the clinical programmer, based on an industry standardplatform, such as personal digital assistant (PDA) or pocket PC, mightinclude the ability to connect to the clinician's computer system in hisor hers office. Such features may take advantage of the PDA systemsoftware for network communications. Such a connection then wouldtransfer relevant patient data to the host computer or server forelectronic processing and archiving. With a feature as described here,the clinical programmer then becomes an integral link in an electronicchain that provides better patient service by reducing the amount ofpaperwork that the physician's office needs to process on each officevisit. It also improves the reliability of the service since it reducesthe chance of mis-entered or misplaced information, such as the recordof the parameter setting adjusted during the visit.

With the use of either the implant charger controller 102, or a patientcontroller 114 (see FIG. 23C), the wireless link 112 allows a patient tocontrol certain predefined parameters of the implantable pulse generatorwithin a predefined limited range. The parameters may include theoperating modes/states, increasing/decreasing or optimizing stimuluspatterns, or providing open or closed loop feedback from an externalsensor or control source. The wireless telemetry 112 also desirablyallows the user to interrogate the implantable pulse generator 18 as tothe status of its internal battery 34. The full ranges within whichthese parameters may be adjusted by the user are controlled, adjusted,and limited by a clinician, so the user may not be allowed the fullrange of possible adjustments.

In one embodiment, the patient controller 114 is sized and configured tocouple to a key chain, as seen in FIG. 23C. It is to be appreciated thatthe patient controller 114 may take on any convenient shape, such as aring on a finger, or a watch on a wrist, or an attachment to a belt, forexample. It may also be desirable to separate the functions of theimplant charger controller 102 into a charger and a patient controller.

The wireless telemetry may incorporate a suitable, low power wirelesstelemetry transceiver (radio) chip set that can operate in the MICS(Medical Implant Communications Service) band (402 MHz to 405 MHz) orother VHF/UHF low power, unlicensed bands. A wireless telemetry link notonly makes the task of communicating with the implantable pulsegenerator 18 easier, but it also makes the link suitable for use inmotor control applications where the user issues a command to theimplantable pulse generator to produce muscle contractions to achieve afunctional goal (e.g., to stimulate ankle flexion to aid in the gait ofan individual after a stroke) without requiring a coil or othercomponent taped or placed on the skin over the implanted implantablepulse generator.

Appropriate use of power management techniques is important to theeffective use of wireless telemetry. Desirably, the implantable pulsegenerator is exclusively the communications slave, with allcommunications initiated by the external controller (the communicationsmaster). The receiver chip of the implantable pulse generator is OFFabout 99% or more of the time and is pulsed on periodically to searchfor a command from an external controller, including but not limited tothe clinical programmer 108, the patient controller 114, the networkinterface 116, and the implant charger controller 102. When theimplantable pulse generator 18 operates at a low rate of wirelesstelemetry because of a low battery, the transceiver chip may be pulsedon less frequently, such as about every five seconds to about tenseconds, to search for a command from an external controller.

Communications protocols include appropriate received message integritytesting and message acknowledgment handshaking to assure the necessaryaccuracy and completeness of every message. Some operations (such asreprogramming or changing stimulus parameters) require rigorous messageaccuracy testing and acknowledgement. Other operations, such as a singleuser command value in a string of many consecutive values, might requireless rigorous checking and no acknowledgement or a more loosely coupledacknowledgement.

The timing with which the implantable pulse generator enables itstransceiver to search for RF telemetry from an external controller isprecisely controlled (using a time base established by a quartz crystal)at a relatively low rate, e.g., the implantable pulse generator may lookfor commands from the external controller for about two milliseconds ata rate of two (2) Hz or less. This equates to a monitoring interval ofabout ½ second or less. It is to be appreciated that implantable pulsegenerator's enabled transceiver rate and the monitoring rate may varyfaster or slower depending on the application. This precise timingallows the external controller to synchronize its next command with thetime that the implantable pulse generator will be listening forcommands. This, in turn, allows commands issued within a short time(seconds to minutes) of the last command to be captured and acted uponwithout having to ‘broadcast’ an idle or pause signal for a fullreceived monitoring interval before actually issuing the command inorder to know that the implantable pulse generator will have enabled itsreceiver and be ready to receive the command. Similarly, thecommunications sequence is configured to have the external controllerissue commands in synchronization with the implantable pulse generatorlistening for commands. Similarly, the command set implemented isselected to minimize the number of messages necessary and the length ofeach message consistent with the appropriate level of error detectionand message integrity monitoring. It is to be appreciated that themonitoring rate and level of message integrity monitoring may varyfaster or slower depending on the application, and may vary over timewithin a given application.

A suitable radio chip is used for the half duplex wirelesscommunications, e.g., the AMIS-52100 (AMI Semiconductor; Pocatello,Id.). This transceiver chip is designed specifically for the MICS andits European counter-part the ULP-AMI (Ultra Low Power-Active MedicalImplant) band. This chip set is optimized by micro-power operation withrapid start-up, and RF ‘sniffing’ circuitry.

The implant charger controller 102 and the implantable pulse generator18, as shown in FIGS. 22A and 22B may also use wireless telemetry toprovide a “smart charge” feature to indicate that charging is occurringand to make corrections to allow for optimal recharging and protectagainst overcharging. During a battery recharge period, the smart chargecauses the implant charger controller 102 to issue commands to theimplantable pulse generator 18 at timed intervals, e.g., every thirtyseconds, to instruct the implantable pulse generator to confirm that thegenerated RF magnetic field is being received and is adequate forrecharging the rechargeable battery. If the implant charger controller102 does not receive a response from the implantable pulse generator 18to confirm that the generated RF magnetic field is being received, theimplant charger controller may stop generating the RF magnetic field.

During the battery recharge period, the implantable pulse generator 18will transmit status information, e.g., an indication of the battery 34charge status and an indication of the magnitude of power recovered bythe receive coil 42, back to the implant charger controller 102.

Based on the magnitude of the power recovered, the smart charge allowsthe implant charger controller 102 to automatically adjust up or downthe magnitude of the magnetic field 100 and/or to instruct the user toreposition the charging coil 104 based on the status information toallow optimal recharging of the implantable pulse generator battery 34while minimizing unnecessary power consumption by the implant chargercontroller 102 and power dissipation in the implantable pulse generator18 (through circuit losses and/or through absorption by the implantablepulse generator case 20 and other components). The magnitude of the RFmagnetic field 100 may be automatically adjusted up to about 300 percentor more of the initial magnitude of the RF magnetic field and adjusteddown until the implant charger controller stops generating the RFmagnetic field.

The instructions to the user to reposition the charging coil 104 may bea visual instruction, such as a bar graph on the implant chargercontroller 102, or a display on the implant charger controller showingrelative positions of the charging coil 104 and the implantable pulsegenerator 18, or an audio instruction, such as a varying tone toindicate relative position, or a combination of instructions.

The smart charge allows for the outer surface of the case 20 of theimplantable pulse generator 18 to maintain a two degree Celsius or lesstemperature rise during the time period in which the receive coil 42 istranscutaneously receiving externally generated power, i.e., RF magneticfield.

In cases where two implant charger controllers 102 could be erroneouslyswapped, or where two or more implantable pulse generators 18 may bewithin wireless telemetry range of each other, e.g., when two users livein the same home, a first implantable pulse generator 18 couldcommunicate with its implant charger controller 102 even when thecharging coil 104 is erroneously positioned over another implantablepulse generator 18. The implant charger controller 102 is configured tocommunicate and charge a specifically identified implantable pulsegenerator (identified by the unique signature/serial number). Becausethe first implantable pulse generator, the one communicating with theimplant charger controller 102, does not sense the RF magnetic chargingfield 100 when the charging coil 104 is positioned over anotherimplantable pulse generator, the first implantable pulse generatorcommunicates with the implant charger controller 102 to increase themagnitude of the RF magnetic field 100. This communication may continueuntil the magnitude of the RF magnetic field is at its maximum.

In order to stop an implant charger controller 102 from attempting tocharge the incorrect implantable pulse generator 18, the implant chargercontroller periodically decreases the magnitude of the RF magnetic field100 and communicates with its (identified by the unique signature/serialnumber) implantable pulse generator to confirm/determine that theimplantable pulse generator 18 sensed the decrease in the magnitude. Ifthe charging coil is erroneously positioned over another implantablepulse generator 18, the correct implantable pulse generator will notsense the decrease and will indicate to the implant charger controller102 that it did not sense the decrease. The implant charger controller102 will then restore the original RF magnetic field strength and retrythe reduced RF magnetic field test. Multiple failures of the test willcause the implant charger controller 102 to suspend charging and notifythe user of the error. Similarly, should the implanted pulse generatornot recover usable power from the RF magnetic field 100 after a fewminutes, the implant charger controller 102 will suspend charging andnotify the user of the error.

d. Stimulus Output Stage

According to one desirable technical feature, the implantable pulsegenerator 18 desirably uses a single stimulus output stage 136(generator) that is directed to one or more output channels (electrodesurfaces) by analog switch(es) or analog multiplexer(s). Desirably, theimplantable pulse generator 18 will deliver at least one channel ofstimulation via a lead/electrode. For applications requiring morestimulus channels, several channels (perhaps up to four) can begenerated by a single output stage. In turn, two or more output stagescould be used, each with separate multiplexing to multiple channels, toallow an implantable pulse generator with eight or more stimuluschannels. As a representative example, the stimulation desirably has abiphasic waveform (net DC current less than 10 microAmps), adjustablefrom about 0.5 mA to about 20 mA based on electrode type and the tissuetype being stimulated, and pulse durations adjustable from about 5microseconds or less up to 500 microseconds or more. The stimuluscurrent (amplitude) and pulse duration being programmable on a channelto channel basis and adjustable over time based on a clinicallyprogrammed sequence or regime or based on user (patient) commandsreceived via the wireless communications link.

A circuit diagram showing a desired configuration for the stimulusoutput stage feature is shown in FIG. 27. It is to be appreciated thatmodifications to this circuit diagram configuration which produce thesame or similar functions as described are within the scope of theinvention.

For neuromodulation/central nervous system applications, the implantablepulse generator 18 may have the capability of applying stimulationtwenty-four hours per day. A typical stimulus regime for suchapplications might have a constant stimulus phase, a no stimulus phase,and ramping of stimulus levels between these phases. For FunctionalElectrical Stimulation (FES), the intensity and timing of thestimulation may vary with user inputs via switches or sensors.

Desirably, the implantable pulse generator 18 includes a single stimulusgenerator (with its associated DC current blocking output capacitor)which is multiplexed to a number of output channels; or a small numberof such stimulus generators each being multiplexed to a number of outputchannels. This circuit architecture allows multiple output channels withvery little additional circuitry. A typical, biphasic stimulus pulse isshown in FIG. 28. Note that the stimulus output stage circuitry 136 mayincorporate a mechanism to limit the recovery phase current to a smallvalue (perhaps 0.5 mA). Also note that the stimulus generator (and theassociated timing of control signals generated by the microcontroller)may provide a delay (typically of the order of 100 microseconds) betweenthe cathodic phase and the recovery phase to limit the recovery phasediminution of the cathodic phase effective at eliciting a neuralexcitation. The charge recovery phase for any electrode (cathode) mustbe long enough to assure that all of the charge delivered in thecathodic phase has been returned in the recovery phase, e.g., greaterthan or equal to five time constants are allowed for the recovery phase.This will allow the stimulus stage to be used for the next electrodewhile assuring there is no net DC current transfer to any electrode.Thus, the single stimulus generator having this characteristic would belimited to four channels (electrodes), each with a maximum frequency of30 Hz to 50 Hz. This operating frequency exceeds the needs of manyindications for which the implantable pulse generator is well suited.For applications requiring more channels (or higher composite operatingfrequencies), two or more separate output stages might each bemultiplexed to multiple (e.g., four) electrodes. Alternatively, theoutput multiplexer/switch stage might allow each output channel to haveits own output coupling capacitor.

e. The Lead Connection Header

According to one desirable technical feature, the implantable pulsegenerator 18 desirably includes a lead connection header 26 forconnecting the lead(s) 12 that will enable reliable and easy replacementof the lead/electrode (see FIGS. 3A and 3B), and includes a smallantenna 80 for use with the wireless telemetry feature.

The implantable pulse generator desirably incorporates a connectionheader (top header) 26 having a conventional connector 82 that is easyto use, reliable, and robust enough to allow multiple replacements ofthe implantable pulse generator after many years (e.g., more than tenyears) of use. The surgical complexity of replacing an implantable pulsegenerator is usually low compared to the surgical complexity ofcorrectly placing the implantable lead 12/electrode 16 in proximity tothe target nerve/tissue and routing the lead 12 to the implantable pulsegenerator. Accordingly, the lead 12 and electrode 16 desirably has aservice life of at least ten years with a probable service life offifteen years or more. Based on the clinical application, theimplantable pulse generator may not have this long a service life. Theimplantable pulse generator service life is largely determined by thepower capacity of the Lithium Ion battery 34, and is likely to be threeto ten years, based on the usage of the device. Desirably, theimplantable pulse generator 18 has a service life of at least fiveyears.

As described above, the implantable pulse generator preferably will usea laser welded titanium case. As with other active implantable medicaldevices using this construction, the implantable lead(s) 12 connect tothe implantable pulse generator through the molded or cast polymericconnection header 26. Metal-ceramic or metal-glass feed-thrus 44, 46, 48(see FIGS. 7 and 16), maintain the hermetic seal of the titanium capsulewhile providing electrical contact to the electrical contacts of thelead 12/electrode 16.

The standard implantable connectors may be similar in design andconstruction to the low-profile IS-1 connector system (per ISO 5841-3).The IS-1 connectors have been in use since the late 1980s and have beenshown to be reliable and provide easy release and re-connection overseveral implantable pulse generator replacements during the service lifeof a single pacing lead. Full compatibility with the IS-1 standard, andmating with pacemaker leads, is not a requirement for the implantablepulse generator.

The implantable pulse generator connection system may include amodification of the conventional IS-1 connector system, which shrinksthe axial length dimensions or adds a third or more electrical contact“rings” or “bands” while keeping the general format and radialdimensions of the IS-1. For application with more than two electrodeconductors, the top header 26 may incorporate one or more connectionreceptacles each of which accommodate leads with typically fourconductors. When two or more leads are accommodated by the header, theselead may exit the connection header in the same or opposite directions(i.e., from opposite sides of the header).

These connectors can be similar to the banded axial connectors used byother multi-polar implantable pulse generators or may follow theguidance of the draft IS-4 implantable connector standard. The design ofthe implantable pulse generator case 20 and header 26 preferablyincludes provisions for adding the additional feed-thrus and largerheaders for such indications.

The inclusion of the UHF antenna 80 for the wireless telemetry insidethe connection header 26 is necessary as the shielding offered by thetitanium case will severely limit (effectively eliminate) radio wavepropagation through the case. The antenna 80 connection will be madethrough feed-thru 48 similar to that used for the electrode connections44, 46. Alternatively, the wireless telemetry signal may be coupledinside the implantable pulse generator onto a stimulus output channeland coupled to the antenna 80 with passive filtering/couplingelements/methods in the connection header 26.

f. The Microcontroller

According to one desirable technical feature, the implantable pulsegenerator 18 desirably uses a standard, commercially availablemicro-power, flash (in-circuit programmable) programmablemicrocontroller 36 or processor core in an application specificintegrated circuit (ASIC). This device (or possibly more than one suchdevice for a computationally complex application with sensor inputprocessing) and other large semiconductor components may have custompackaging such as chip-on-board, solder flip chip, or adhesive flip chipto reduce circuit board real estate needs.

A circuit diagram showing a desired configuration for themicrocontroller 36 is shown in FIG. 29. It is to be appreciated thatmodifications to this circuit diagram configuration which produce thesame or similar functions as described are within the scope of theinvention.

g. Power Management Circuitry

According to one desirable technical feature, the implantable pulsegenerator 18 desirably includes efficient power management circuitry asan element of the implantable pulse generator circuitry 32 shown in FIG.24. The power management circuitry is generally responsible for theefficient distribution of power and monitoring the battery 34, and forthe recovery of power from the RF magnetic field 100 and for chargingand monitoring the battery 34. In addition, the operation of theimplantable pulse generator 18 can be described in terms of havingoperating modes as relating to the function of the power managementcircuitry. These modes may include, but are not limited to IPG Active,IPG Dormant, and, IPG Active and Charging. These modes will be describedbelow in terms of the principles of operation of the power managementcircuitry using possible circuit diagrams shown in FIGS. 30 and 31. FIG.30 shows one possible power management sub-circuit having MOSFETisolation between the battery 34 and the charger circuit. FIG. 31 showsanother possible power management sub-circuit diagram without havingMOSFET isolation between the battery 34 and the charger circuit. In thecircuit without the isolation MOSFET (see FIG. 31), the leakage currentof the disabled charge control integrated circuit chip (U1) must be verylow to prevent this leakage current from discharging the battery 34 inall modes (including the Dormant mode). Except as noted, the descriptionof these modes applies to both circuits.

i. IPG Active Mode

The IPG Active mode occurs when the implantable pulse generator 18 isoperating normally. In this mode, the implantable pulse generator may begenerating stimulus outputs or it may be waiting to generate stimulus inresponse to a timed neuromodulation sequence or a telemetry command froman external controller. In this mode, the implantable pulse generator isactive (microcontroller 36 is powered and coordinating wirelesscommunications and may be timing & controlling the generation anddelivery of stimulus pulses).

i(a) Principles of Operation, IPG Active Mode

In the IPG Active mode, the lack of a RF magnetic field from a chargingcoil means there will be no DC current from VRAW, which means that Q5 isheld off (see FIG. 30). This, in turn, holds Q3 off and a portion of thepower management circuitry is isolated from the battery 34. In FIG. 31,the lack of DC current from VRAW means that U1 is disabled eitherdirectly or via the microcontroller. This, in turn, keeps the currentdrain from the battery 34 to an acceptably low level, typically lessthan one microAmp.

ii. IPG Dormant Mode

The IPG Dormant mode occurs when the implantable pulse generator 18 iscompletely disabled (powered down). In this mode, power is not beingsupplied to the microcontroller 36 or other enabled circuitry. This isthe mode for the long-term storage of the implantable pulse generatorbefore or after implantation. As a safety feature, the Dormant mode mayalso be entered by placing a pacemaker magnet 118 (or comparable device)over the implantable pulse generator 18 for a predetermined amount oftime, e.g., five seconds. The implantable pulse generator 18 may also beput in the Dormant mode by a wireless telemetry command from an externalcontroller.

The Dormant mode may be exited by placing the implantable pulsegenerator 18 into the Active and Charging mode by placing the chargingcoil 104 of a functional implant charger controller 102 in closeproximity to the implantable pulse generator 18.

ii(a) Principles of Operation, IPG Dormant Mode

In the IPG Dormant mode, VBAT is not delivered to the remainder of theimplantable pulse generator circuitry because Q4 is turned off. TheDormant mode is stable because the lack of VBAT means that VCC is alsonot present, and thus Q6 is not held on through R8 and R10. Thus thebattery 34 is completely isolated from all load circuitry (the VCC powersupply and the VHH power supply).

The Dormant mode may be entered through the application of the magnet118 placement over S1 (magnetic reed switch) or through the reception ofa command by the wireless telemetry. In the case of the telemetrycommand, the PortD4, which is normally configured as a microcontrollerinput, is configured as a logic output with a logic low (0) value. This,in turn, discharges C8 through R12 and turns off Q6; which, in turn,turns off Q4 and forces the implantable pulse generator into the Dormantmode. Note that R12 is much smaller in value than R10, thus themicrocontroller 36 can force C8 to discharge even though VCC is stillpresent.

In FIG. 30, the lack of DC current from VRAW means that Q5 is held off.This, in turn, holds Q3 off and a portion of the power managementcircuitry is isolated from the battery 34. Also, Q4 was turned off. InFIG. 13, the lack of DC current from VRAW means that U1 is disabled.This, in turn, keeps the current drain from the battery 34 to anacceptably low level, typically less than 1 μA.

iii. IPG Active and Charging Mode

In the embodiment having a rechargeable battery, the implantable pulsegenerator Active and Charging mode occurs when the implantable pulsegenerator 18 is being charged. In this mode, the implantable pulsegenerator 18 is active, i.e., the microcontroller 36 is powered andcoordinating wireless communications and may be timing and controllingthe generation and delivery of stimulus pulses. The implantable pulsegenerator 18 may use the smart charge feature to communicate with theimplant charger controller 102 concerning the magnitude of the batteryvoltage and the DC voltage recovered from the RF magnetic field 100. Theimplant charger controller 102 uses this data for two purposes: toprovide feedback to the user about the proximity of the charging coil104 to the implanted pulse generator, and to increase or decrease thestrength of the RF magnetic field 100. This, in turn, minimizes thepower losses and undesirable heating of the implantable pulse generator.

While in the IPG Active and Charging mode, the power managementcircuitry 130 serves the following primary functions:

(1) provides battery power to the rest of the implantable pulsegenerator circuitry 32,

(2) recovers power from the RF magnetic field 100 generated by theimplant charger controller 102,

(3) provides controlled charging current (from the recovered power) tothe battery 34, and

(4) communicates with the implant charger controller 102 via thewireless telemetry link 112 to provide feedback to the user positioningthe charging coil 104 over the implantable pulse generator 18, and tocause the implant charger controller 102 to increase or decrease thestrength of its RF magnetic field 100 for optimal charging of theimplantable pulse generator battery 34 (Lithium Ion battery).

iii(a) Principles of Operation, IPG Active and Charging Mode

-   -   iii(a)(1) RF voltage is induced in the receive coil 42 by the RF        magnetic field 100 of the implant charger controller 102    -   iii(a)(2) Capacitor C1 is in series with the receive coil and is        selected to introduce a capacitive reactance that compensates        (subtracts) the inductive reactance of the receive coil 42    -   iii(a)(3) D1-D2 form a full wave rectifier that converts the AC        voltage recovered by the receive coil 42 into a pulsating DC        current flow    -   iii(a)(4) This pulsating DC current is smoothed (filtered) by C3        (this filtered DC voltage is labeled VRAW)    -   iii(a)(5) D4 is a zener diode that acts as a voltage limiting        device (in normal operation, D4 is not conducting significant        current)    -   iii(a)(6) D3 prevents the flow of current from the battery 34        from preventing the correct operation of the power management        circuitry 130 once the voltage recovered from the RF magnetic        field is removed. Specifically, current flow from the battery        [through Q3 (turned ON), in the case for the circuit of FIG. 30]        through the body diode of Q2 would hold ON the charge controller        IC (U1). This additional current drain would be present in all        modes, including Dormant, and would seriously limit battery        operating life. Additionally, this battery current pathway would        keep Q6 turned ON even if the magnetic reed switch (S1) were        closed; thus preventing the isolation of the implantable pulse        generator circuitry from the battery in the Dormant mode.    -   iii(a)(7) U1, Q2, R2, C4, C6, and C2 form the battery charger        sub-circuit        -   U1 is a micropower, Lithium Ion Charge Management Controller            chip implementing a constant current phase and constant            voltage phase charge regime. This chip desirably            incorporates an internal voltage reference of high accuracy            (+/−0.5%) to establish the constant voltage charge level. U1            performs the following functions:        -   monitors the voltage drop across a series resistor R2            (effectively the current charging the battery 34) to control            the current delivered to the battery through the external            pass transistor Q2. U1 uses this voltage across R2 to            establish the current of the constant current phase            (typically the battery capacity divided by five hours) and        -   decreases the current charging the battery as required to            limit the battery voltage and effectively transition from            constant current phase to constant voltage phase as the            battery voltage approaches the terminal voltage,    -   iii(a)(8) U1 may also include provisions for timing the duration        of the constant current and constant voltage phases and suspends        the application of current to the battery 34 if too much time is        spent in the phase. These fault timing features of U1 are not        used in normal operation.    -   iii(a)(9) In this circuit, the constant voltage phase of the        battery charging sequence is timed by the microcontroller 36 and        not by U1. The microcontroller monitors the battery voltage and        terminates the charging sequence (i.e., tells the implant        charger controller 102 that the implantable pulse generator        battery 34 is fully charged) after the battery voltage has been        in the constant voltage region for greater than a fixed time        period (e.g., 15 to 20 minutes).    -   iii(a)(10) In FIGS. 30, Q3 and Q5 are turned ON only when the        charging power is present. This effectively isolates the        charging circuit from the battery 34 when the externally        supplied RF magnetic field 100 is not present and providing        power to charge the rechargeable battery.    -   iii(a)(11) In FIG. 31, U1 is always connected to the battery 34,        and the disabled current of this chip is a load on the battery        34 in all modes (including the Dormant mode). This, in turn, is        a more demanding requirement on the current consumed by U1 while        disabled.    -   iii(a)(12) F1 is a fuse that protects against long-duration,        high current component failures. In most transient high current        failures, (i.e., soft failures that cause the circuitry to        consume high current levels and thus dissipate high power        levels; but the failure initiating the high current flow is not        permanent and the circuit will resume normal function if the        circuit is removed from the power source before damage from        overheating occurs), the VBAT circuitry will disconnect the        battery 34 from the temporary high load without blowing the        fuse. The specific sequence is:        -   High current flows into a component(s) powered by VBAT (most            likely the VHH power supply or an element powered by the VCC            power supply).        -   The voltage drop across the fuse will (prior to the fuse            blowing) turn ON Q1 (based on the current flow through the            fuse causing a 0.5V to 0.6V drop across the resistance of            F1).        -   The collector current from Q1 will turn off Q4.        -   VBAT drops very quickly and, as a direct result, VCC falls.            In turn, the voltage on the PortD4 IO pin from the            microcontroller voltage falls as VCC falls, through the            parasitic diodes in the microcontroller 36. This then pulls            down the voltage across C6 as it is discharged through R12.        -   The implantable pulse generator 18 is now stable in the            Dormant mode, i.e., VBAT is disconnected from the battery 34            by a turned OFF Q4. The only load remaining on the battery            is presented by the leakage current of the charging circuit            and by the analog multiplexer (switches) U3 that are used to            direct an analog voltage to the microcontroller 36 for            monitoring the voltage after the resistance of F1) an            estimate of the current consumption of the entire circuit. A            failure of these voltage monitoring circuits is not            protected by the fuse, but resistance values limit the            current flow to safe levels even in the event of component            failures. A possible source of a transient high-current            circuit failure is the SCR latchup or supply-to-ground short            failure of a semiconductor device directly connected to VBAT            or VCC.    -   iii(a)(13) R9 & R11 form a voltage divider to convert VRAW (0V        to 8V) into the voltage range of the microcontroller's A-D        inputs (used for closed loop control of the RF magnetic field        strength),    -   iii(a)(14) R8 and C9 form the usual R-C reset input circuit for        the microcontroller 36; this circuit causes a hardware reset        when the magnetic reed switch (S1) is closed by the application        of a suitable static magnetic field for a short duration,    -   iii(a)(15) R10 and C8 form a much slower time constant that        allows the closure of the reed switch by the application of the        static magnetic field for a long duration to force the        implantable pulse generator 18 into the Dormant mode by turning        OFF Q6 and thus turning OFF Q4. The use of the magnetic reed        switch for resetting the microcontroller 36 or forcing a total        implantable pulse generator shutdown (Dormant mode) may be a        desirable safety feature.

2. Representative Implant Pulse

Generator Circuitry

FIG. 24 shows an embodiment of a block diagram circuit 32 for therechargeable implantable pulse generator 18 that takes into account thedesirable technical features discussed above. FIG. 25 shows anembodiment of a block diagram circuit 33 for the implantable pulsegenerator 88 that also takes into account the desirable technicalfeatures discussed above.

Both the circuit 32 and the circuit 33 can be grouped into functionalblocks, which generally correspond to the association andinterconnection of the electronic components. FIGS. 24 and 25 showalternative embodiments of a block diagram that provides an overview ofa representative desirable implantable pulse generator design. As can beseen, there may be re-use of the circuit 32, or alternatively, portionsof the circuit 32 of the rechargeable implantable pulse generator 18,with minimal modifications, e.g., a predetermined selection ofcomponents may be included or may be exchanged for other components, andminimal changes to the system operating software (firmware). Re-use of amajority of the circuitry from the rechargeable implantable pulsegenerator 18 and much of the firmware allows for a low development costfor the rechargeable and primary cell implantable pulse generator.

In FIGS. 24 and 25, seven functional blocks are shown: (1) TheMicroprocessor or Microcontroller 36; (2) the Power Management Circuit130; (3) the VCC Power Supply 132; (4) the VHH Power Supply 134; (5) theStimulus Output Stage(s) 136; (6) the Output Multiplexer(s) 138; and (7)the Wireless Telemetry Circuit 140.

For each of these blocks, the associated functions, possible keycomponents, and circuit description are now described.

a. The Microcontroller

The Microcontroller 36 is responsible for the following functions:

(1) The timing and sequencing of the stimulus output stage 136 and theVHH power supply 134 used by the stimulus output stage,

(2) The sequencing and timing of power management functions,

(3) The monitoring of the battery voltage, the stimulator voltagesproduced during the generation of stimulus pulses, and the total circuitcurrent consumption, VHH, and VCC,

(4) The timing, control, and interpretation of commands to and from thewireless telemetry circuit 140,

(5) The logging (recording) of patient usage data as well as clinicianprogrammed stimulus parameters and configuration data, and

(6) The processing of commands (data) received from the user (patient)via the wireless link to modify the characteristics of the stimulusbeing delivered or to retrieve logged usage data.

The use of a microcontroller incorporating flash programmable memoryallows the operating system software of the implantable pulse generatoras well as the stimulus parameters and settings to be stored innon-volatile memory (data remains safely stored even if the battery 34becomes fully discharged; or if the implantable pulse generator isplaced in the Dormant mode). Yet, the data (operating system software,stimulus parameters, usage history log, etc.) can be erased andreprogrammed thousands of times during the life of the implantable pulsegenerator. The software (firmware) of the implantable pulse generatormust be segmented to support the operation of the wireless telemetryroutines while the flash memory of the microcontroller 36 is beingerased and reprogrammed. Similarly, the VCC power supply 132 mustsupport the power requirements of the microcontroller 36 during theflash memory erase and program operations.

Although the microcontroller 36 may be a single component, the firmwareis developed as a number of separate modules that deal with specificneeds and hardware peripherals. The functions and routines of thesesoftware modules are executed sequentially; but the execution of thesemodules are timed and coordinated so as to effectively functionsimultaneously. The microcontroller operations that are associateddirectly with a given hardware functional block are described with thatblock.

The Components of the Microcontroller Circuit may include:

-   -   (1) A single chip microcontroller 36. This component may be a        member of the Texas Instruments MSP430 family of flash        programmable, micro-power, highly integrated mixed signal        microcontroller. Likely family members to be used include the        MSP430F1610, MSP430F1611, MSP430F1612, MSP430F168, and the        MSP430F169. Each of these parts has numerous internal        peripherals, and a micropower internal organization that allows        unused peripherals to be configured by minimal power        dissipation, and an instruction set that supports bursts of        operation separated by intervals of sleep where the        microcontroller suspends most functions.    -   (2) A miniature, quartz crystal (X1) for establishing precise        timing of the microcontroller. This may be a 32.768 KHz quartz        crystal.    -   (3) Miscellaneous power decoupling and analog signal filtering        capacitors.

b. Power Management Circuit

The Power Management Circuit 130 (including associated microcontrolleractions) is responsible for the following functions:

(1) monitor the battery voltage,

(2) suspend stimulation when the battery voltage becomes very low,and/or suspend all operation (go into the Dormant mode) when the batteryvoltage becomes critically low,

(3) communicate (through the wireless telemetry link 112) with theexternal equipment the charge status of the battery 34,

(4) prevent (with single fault tolerance) the delivery of excessivecurrent from the battery 34,

(5) provide battery power to the rest of the circuitry of theimplantable pulse generator, e.g., VCC and VHH power supplies,

(6) provide isolation of the Lithium Ion battery 34 from other circuitrywhile in the Dormant mode,

(7) provide a hard microprocessor reset and force the implantable pulsegenerator 18 into the Dormant mode in the presence of long pacemakermagnet 118 application (or comparable device),

(8) provide the microcontroller 36 with analog voltages with which tomeasure the magnitude of the battery voltage and the appropriate batterycurrent flow (drain and charge),

(9) recover power from the receive coil 42,

(10) control delivery of the receive coil power to recharge the LithiumIon battery 34,

(11) monitor the battery voltage during charge and discharge conditions,

(12) communicate (through the wireless telemetry link 112) with theexternally mounted or worn implant charger controller 102 to increase ordecrease the strength of the RF magnetic field 100 for optimal chargingof the battery 34,

(13) disable (with single fault tolerance) the delivery of chargingcurrent to the battery 34 in overcharge conditions, and

(14) provide the microcontroller 36 with analog voltages with which tomeasure the magnitude of the recovered power from the RF magnetic field100.

The Components of the Power Management Circuit may include:

(1) Low on resistance, low threshold P channel MOSFETs with very low offstate leakage current (Q2, Q3, and Q4).

(2) Analog switches (or an analog multiplexer) U3.

(3) Logic translation N-channel MOSFETs (Q5 & Q6) with very low offstate leakage current.

(4) The receive coil 42 (see FIGS. 4B, 4C, and 4D), which desirably is amulti-turn, fine copper wire coil near the inside perimeter of theimplantable pulse generator 18. Preferably, the receive coil includes apredetermined construction, e.g., 300 turns, each of four strands of #40enameled magnetic wire, or the like. The maximizing of the coil'sdiameter and reduction of its effective RF resistance allows necessarypower transfer at and beyond the typical implant depth of about onecentimeter.

As can be seen in FIG. 4C, the receive coil 42 is generally rectangularin cross sectional shape, with a height H greater than its width W. Inone embodiment, the height H is about five millimeters to about sixmillimeters, and the width W is about two millimeters to threemillimeters.

The receive coil 42 also includes a maximum outside dimension X of aboutseventeen millimeters to about twenty millimeters, for example, as shownin FIG. 4D. The maximum outside dimension X may be measured from themidpoint on a straight line that bisects the coil into two equal parts.Although there may be more than one line that bisects the coil 42, thedimension X is to be the longest dimension X possible from the midpointof the bisection line to the coil's outside edge.

(5) A micropower Lithium Ion battery charge management controller IC(integrated circuit); such as the MicroChip MCP73843-41, or the like.

c. The VCC Power Supply

The VCC Power Supply 132 is generally responsible for the followingfunctions:

(1) Some of the time, the VCC power supply passes the battery voltage tothe circuitry powered by VCC, such as the microcontroller 36, stimulusoutput stage 136, wireless telemetry circuitry 140, etc.

(2) At other times, the VCC power supply fractionally steps up thevoltage to about 3.3V; (other useable voltages include 3.0V, 2.7V, etc.)despite changes in the Lithium Ion battery 34 voltage. This highervoltage is required for some operations such as programming or erasingthe flash memory in the microcontroller 36, (e.g., in circuitprogramming).

The voltage converter/switch part at the center of the VCC power supplymay be a charge pump DC to DC converter. Typical choices for this partmay include the Maxim MAX1759, the Texas Instruments TPS60204, or theTexas Instruments REG710, among others. In the embodiment having arechargeable battery 34, the VCC power supply may include a micropower,low drop out, linear voltage regulator; e.g., Microchip NCP1700T-3302,Maxim Semiconductor MAX1725, or Texas Instruments TPS79730.

The characteristics of the VCC Power Supply might include:

(1) high efficiency and low quiescent current, i.e., the current wastedby the power supply in its normal operation. This value should be lessthan a few microamperes; and

(2) drop-out voltage, i.e., the minimal difference between the VBATsupplied to the VCC power supply and its output voltage. This voltagemay be less than about 100 mV even at the current loads presented by thetransmitter of the wireless telemetry circuitry 140.

(3) The VCC power supply 132 may allow in-circuit reprogramming of theimplantable pulse generator firmware.

d. VHH Power Supply

A circuit diagram showing a desired configuration for the VHH powersupply 134 is shown in FIG. 32. It is to be appreciated thatmodifications to this circuit diagram configuration which produce thesame or similar functions as described are within the scope of theinvention. The VHH power supply 134 is generally responsible for thefollowing functions:

(1) Provide the Stimulus Output Stage 136 and the Output Multiplexer 138with a programmable DC voltage between the battery voltage and a voltagehigh enough to drive the required cathodic phase current through theelectrode circuit plus the voltage drops across the stimulator stage,the output multiplexer stage, and the output coupling capacitor. VHH istypically 12 VDC or less for neuromodulation applications; and 25V orless for muscle stimulation applications, although it may be higher forvery long lead lengths.

The Components of the VHH Power Supply might include:

(1) Micropower, inductor based (fly-back topology) switch mode powersupply (U10); e.g., Texas Instruments TPS61045, Texas InstrumentsTPS61041, or Linear Technology LT3464 with external voltage adjustmentcomponents.

(2) L6 is the flyback energy storage inductor.

(3) C42 & C43 form the output capacitor.

(4) R27, R28, and R29 establish the operating voltage range for VHHgiven the internal DAC which is programmed via the SETVHH logic commandfrom the microcontroller 36.

(5) Diode D9 serves no purpose in normal operation and is added to offerprotection from over-voltage in the event of a VHH circuit failure.

(6) The microcontroller 36 monitors VHH for detection of a VHH powersupply failure, system failures, and optimizing VHH for the exhibitedelectrode circuit impedance.

e. Stimulus Output Stage

The Stimulus Output Stage(s) 136 is generally responsible for thefollowing functions:

(1) Generate the identified biphasic stimulus current with programmable(dynamically adjustable during use) cathodic phase amplitude, pulsewidth, and frequency. The recovery phase may incorporate a maximumcurrent limit; and there may be a delay time (most likely a fixed delay)between the cathodic phase and the recovery phase (see FIG. 28). Typicalcurrents (cathodic phase) vary from about 0.5 mA to about 20 mA based onthe electrode construction and the nature of the tissue beingstimulated. Electrode circuit impedances can vary with the electrode andthe application, but are likely to be less than 2,000 ohms and greaterthan 100 ohms across a range of electrode types.

The Components of the Stimulus Output Stage may include:

(1) The cathodic phase current through the electrode circuit isestablished by a high gain (HFE) NPN transistor (Q7) with emitterdegeneration. In this configuration, the collector current of thetransistor (Q7) is defined by the base drive voltage and the value ofthe emitter resistor (R24).

Two separate configurations are possible: In the first configuration (asshown in FIG. 27), the base drive voltage is provided by a DACperipheral inside the microcontroller 36 and is switched on and off by atimer peripheral inside the microcontroller. This switching function isperformed by an analog switch (U8). In this configuration, the emitterresistor (R24) is fixed in value and fixed to ground.

In a second alternative configuration, the base drive voltage is a fixedvoltage pulse (e.g., a logic level pulse) and the emitter resistor ismanipulated under microcontroller control. Typical options may includeresistor(s) terminated by microcontroller IO port pins that are held orpulsed low, high, or floating; or an external MOSFET that pulls one ormore resistors from the emitter to ground under program control. Notethat the pulse timing need only be applied to the base drive logic; thetiming of the emitter resistor manipulation is not critical.

The transistor (Q7) desirably is suitable for operation with VHH on thecollector. The cathodic phase current through the electrode circuit isestablished by the voltage drop across the emitter resistor. Diode D7,if used, provides a degree of temperature compensation to this circuit.

(2) The microcontroller 36 (preferably including a programmablecounter/timer peripheral) generates stimulus pulse timing to generatethe cathodic and recovery phases and the interphase delay. Themicrocontroller 36 also monitors the cathode voltage to confirm thecorrect operation of the output coupling capacitor, to detect systemfailures, and to optimize VHH for the exhibited electrode circuitimpedance; i.e., to measure the electrode circuit impedance.Additionally, the microcontroller 36 can also monitor the pulsingvoltage on the emitter resistor; this allows the fine adjustment of lowstimulus currents (cathodic phase amplitude) through changes to the DACvalue.

f. The Output Multiplexer

The Output Multiplexer 138 is generally responsible for the followingfunctions:

(1) Route the Anode and Cathode connections of the Stimulus Output Stage136 to the appropriate electrode based on addressing data provided bythe microcontroller 36.

(2) Allow recharge (recovery phase) current to flow from the outputcoupling capacitor C36 back through the electrode circuit with aprogrammable delay between the end of the cathodic phase and thebeginning of the recovery phase (the interphase delay).

The circuit shown in FIG. 27 may be configured to provide monopolarstimulation (using the case 20 as the return electrode) to Electrode 1,to Electrode 2, or to both at the same time (sharing the current), orseparately—perhaps with different stimulus parameters —through timemultiplexing. This circuit could also be configured as a single bipolaroutput channel by changing the hardwire connection between the circuitboard 32 and the electrode; i.e., by routing the case 20 connection toElectrode 1 or Electrode 2. The use of four or more channels permultiplexer stage (i.e., per output coupling capacitor) is possible.

The Components of the Output Multiplexer might include:

(1) An output coupling capacitor in series with the electrode circuit.This capacitor is desirably located such that there is no DC across thecapacitor in steady state. This capacitor is typically charged by thecurrent flow during the cathodic phase to a voltage range of about ¼thto 1/10th of the voltage across the electrode circuit during thecathodic phase. Similarly, this capacitor is desirably located in thecircuit such that the analog switches do not experience voltages beyondtheir ground of power supply (VHH).

(2) The analog switches (U7) must have a suitably high operatingvoltage, low ON resistance, and very low quiescent current consumptionwhile being driven from the specified logic levels. Suitable analogswitches include the Vishay/Siliconix DG412HS, for example.

(3) Microcontroller 36 selects the electrode connections to carry thestimulus current (and time the interphase delay) via address lines.

(4) Other, analog switches (U9) may be used to sample the voltage of VHH134, the case 20, and the selected Electrode. The switched voltage,after the voltage divider formed by R25 and R26, is monitored by themicrocontroller 36.

g. Wireless Telemetry Circuit

The Wireless Telemetry circuit 140 is responsible for the followingfunctions:

(1) Provide reliable, bidirectional communications (half duplex) with anexternal controller—e.g., clinical programmer 108 or a implant chargercontroller 102, for example, via a VHF-UHF RF link (likely in the 402MHZ to 405 MHz MICS band per FCC 47 CFR Part 95 and the Ultra LowPower—Active Medical Implant (AMI) regulations of the European Union).This circuit will look for RF commands at precisely timed intervals(e.g., twice a second), and this function must consume very littlepower. Much less frequently this circuit will transmit responses tocommands sent by the external controller. This function should also beas low power as possible, but will represent a lower total energy demandthan the receiver in most of the anticipated applications becausewireless telemetry transmissions by the implantable pulse generator 18will typically be rare events. The RF carrier is amplitude modulated(on-off keyed) with the digital data. Serial data is generated by themicrocontroller 36 already formatted and timed. The wireless telemetrycircuit 140 converts the serial data stream into a pulsing carriersignal during the transmit process; and it converts a varying RF signalstrength into a serial data stream during the receive process (see FIG.26B).

The Components of the Wireless Telemetry Circuit might include:

(1) a crystal controlled, micropower transceiver chip such as the AMISemiconductor AMIS-52100 (U6). This chip is responsible for generatingthe RF carrier during transmissions and for amplifying, receiving, anddetecting (converting to a logic level) the received RF signals. Thetransceiver chip must also be capable of quickly starting and stoppingoperation to minimize power consumption by keeping the chip disabled(and consuming very little power) the majority of the time; and poweringup for only as long as required for the transmitting or receivingpurpose. The transceiver chip may be enabled only when the stimulusoutput stage is not generating stimulus current.

(2) The transceiver chip has separate transmit and receive ports thatmust be switched to a single antenna/feedthru. This function isperformed by the transmit/receive switch (U5) under microcontrollercontrol via the logic line XMIT. The microcontroller 36 controls theoperation of the transceiver chip via an I²C (2-wire serial interface)serial communications link. The serial data to and from the transceiverchip may be handled by a UART or a SPI peripheral of themicrocontroller. The message encoding/decoding and error detection maybe performed by a separate, dedicated microcontroller; else thisprocessing will be time shared with the other tasks of the onlymicrocontroller.

The various inductor and capacitor components (U6) surrounding thetransceiver chip and the transmit/receive switch (U5) are impedancematching components and harmonic filtering components, except asfollows:

(1) X2, C33, and C34 are used to generate the crystal controlledreference frequency, desirably tuned to 1/32 of the desired RF carrierfrequency,

(2) L4 and C27 form the tuned elements of a VCO (voltage controlledoscillator) operating at twice the carrier frequency, and

(3) R20, C29, and C30 are filter components of the PLL (phase lockedloop) filter used to generate the carrier (transmitter) or localoscillator (receiver) frequencies from the reference frequency.

B. Lead and Electrode

As previously described, the system 10 includes an implantable pulsegenerator 18, a lead 12, and an electrode 16. Two possible types ofelectrodes will be described below, although any number of electrodetypes may be used.

In one embodiment, the lead 12 and electrode 16 are sized and configuredto be inserted into and to rest in tissue (see FIG. 2A), such as in thelower abdomen for example, without causing pain or discomfort or impactbody image. Desirably, the lead 12 and electrode 16 can be insertedusing a small (e.g., smaller than 16 gauge) introducer 158 (see FIG. 36)with minimal tissue trauma. The lead 12 and electrode 16 are formed froma biocompatible and electrochemically suitable material and possess nosharp features that can irritate tissue during extended use.Furthermore, the lead 12 and electrode 16 possess mechanicalcharacteristics including mechanical compliance (flexibility) toflexibly respond to dynamic stretching, bending, and crushing forcesthat can be encountered within tissue in a wide variety of body regionswithout damage or breakage, and to accommodate relative movement of thepulse generator 18 coupled to the lead 12 without imposing force ortorque to the electrode 16 which tends to dislodge the electrode.

Furthermore, the lead 12 and electrode 16 desirably include an anchoringmeans 150 for providing retention strength to resist migration within orextrusion from tissue in response to force conditions normallyencountered during periods of extended use (see FIG. 33). In addition,the anchoring means 150 is desirably sized and configured to permit theelectrode 16 position to be adjusted easily during insertion, allowingplacement at the optimal location where selective stimulation may occur.The anchoring means 150 functions to hold the electrode at the implantedlocation despite the motion of the tissue and small forces transmittedby the lead 12 due to relative motion of the coupled implantable pulsegenerator 18 due to changes in body posture or external forces appliedto the implant region. However, the anchoring means 150 should allowreliable release of the electrode 16 at higher force levels, to permitwithdrawal of the implanted electrode 16 by purposeful pulling on thelead 12 at such higher force levels, without breaking or leavingfragments, should removal of the implanted electrode 16 be desired.

The lead 12 and electrode 16 is sized and configured to be anchored insoft adipose tissue, with no dependence on support or stability frommuscle tissue. The lead 12 and electrode 16 are particularly well suitedfor placement in this soft adipose tissue because of the unique shape,size, spacing, and orientation of the anchoring means 150, which allowsthe lead 12 and electrode 16 to be used for other indications, such asin the field of urology (e.g., stimulation of nerves in adipose tissuefor the treatment of incontinence and/or sexual restoration).

1. The Lead

FIG. 33 shows a representative embodiment of a lead 12 and electrode 16that provide the foregoing features. The implantable lead 12 comprises amolded or extruded component 152, which may encapsulate or enclose (inthe case of a tubular construction) a coiled stranded wire element 154,and a plug or connector 155 (shown in FIG. 33). The lead 12 may becomposed of one wire 154 connecting a single electrode 16 to contact(s)of the connector 155. Alternatively, the lead 12 may be composed ofseveral individually insulated wires 154 connecting multiple electrodes16 to multiple contacts of the connector 155. Each wire may be a singlestrand of metal, such as MP35N nickel-cobalt, or 316L stainless steel,or a more complex structure such as drawn tube of MP35N or 316L filledwith silver. Alternatively, each separate insulated wire may be composedof multiple strands of wire (three such strands are shown in FIG. 34A),with each strand electrically connected in parallel at the electrode endand at the connector end. Examples of suitable electrical insulationinclude polyimide, parylene, and polyurethane. The molded or extrudedlead 12 can have an outside diameter as small as about one (1) mm. Thelead 12 may be approximately 10 cm to 40 cm in length, although lengthsextending the length of the body are possible. The lead 12 provideselectrical continuity between the connector 155 and the electrode 16.

The coil's pitch can be constant, as FIG. 34B shows, or the coil's pitchcan alternate from high to low spacing to allow for flexibility in bothcompression and tension, as FIG. 34A shows. The tight pitch will allowfor movement in tension, while the open pitch will allow for movement incompression.

A standard IS-1 or similar type connector 155 at the proximal endprovides electrical continuity and mechanical attachment to theimplantable pulse generator's connector jack 82. The lead 12 andconnector 155 all may include provisions for a guidewire that passesthrough these components and the length of the lead 12 to the conductiveelectrode 16 at the distal end. Such a guidewire or stylet would allowthe easy deployment of the lead 12 through an introducer.

2. The Electrode

The electrode 16 may comprise one or more electrically conductivesurfaces. Two conductive surfaces are show in FIG. 33. The twoconductive surfaces can be used either A) as two individual stimulating(cathodic) electrodes in monopolar configuration using the case 20 ofthe implantable pulse generator 18 as the return (anodic) electrode orB) in bipolar configuration with one electrode functioning as thestimulating (cathodic) electrode and the other as the return (anodic)electrode.

In general, bipolar stimulation is more spatially specific thanmonopolar stimulation—the area of stimulation is much smaller—which isgood if the electrode 16 is close to a targeted tissue region, e.g., anerve. But if the electrode 16 is farther from the target tissue region,then a monopolar configuration could be used because with theimplantable pulse generator 18 acting as the return electrode,activation of the tissue is less sensitive to exact placement than witha bipolar configuration.

Often in use, a physician may first attempt to place the electrode 16close to the target tissue region so that it could be used in a bipolarconfiguration, but if bipolar stimulation failed to activate the targettissue region, then the electrode 16 could be switched to a monopolarconfiguration. Two separate conductive surfaces on the electrode 16provide an advantage because if one conductive surface fails to activatethe target tissue region because it is too far from the target tissueregion, then stimulation with the second conductive surface could betried, which might be closer to the target tissue region. Without thesecond conductive surface, a physician would have to reposition theelectrode to try to get closer to the target tissue region. This sameconcept may be extended to more than two conductive surfaces as well.

The electrode 16, or electrically conductive surface or surfaces, can beformed from PtIr (platinum-iridium) or, alternatively, 316L stainlesssteel or titanium, and possess a conductive surface of approximately 10mm² to 20 mm². This surface area provides current densities up to 2mA/mm² with per pulse charge densities less than 0.5 μC/mm². Thesedimensions and materials deliver a charge safely within the stimulationlevels supplied by the implantable pulse generator.

Each conductive surface has an axial length in the range of about onemillimeter to about five millimeters in length. When two or moreconductive surfaces are used, either in the monopolar or bipolarconfigurations as described, there will be an axial spacing between theconductive surfaces in the range of about one millimeter to about tenmillimeters separation.

3. The Anchoring Means

In the illustrated embodiment (see FIG. 33), the lead is anchored byanchoring means 150 specifically designed to secure the electrode 16 ina targeted tissue region, e.g., the layer of adipose tissue, without thesupport of muscle tissue. The anchoring means 150 takes the form of anarray of shovel-like blades or scallops 156 proximal to theproximal-most electrode 16 (although a blade 156 or blades could also beproximal to the distal most electrode 16, or could also be distal to thedistal most electrode 16). The blades 156 desirably present relativelylarge, generally planar surfaces, and are placed in multiple rowsaxially along the lead 12. The blades 156 may also be somewhat arcuateas well, or a combination of arcuate and planar surfaces. A row ofblades 156 comprises two blades 156 spaced 180 degrees apart. The blades156 may have an axial spacing between rows of blades in the range of sixto fourteen millimeters, and each row may be spaced apart 90 degrees.The blades 156 are normally biased toward a radially outward conditioninto tissue. In this condition, the large surface area and orientationof the blades 156 allow the lead 12 to resist dislodgement or migrationof the electrode 16 out of the correct location in the surroundingtissue. In the illustrated embodiment, the blades 156 are biased towarda proximal-pointing orientation, to better resist proximal migration ofthe electrode 16 with lead tension. The blades 156 are desirably madefrom a polymer material, e.g., high durometer silicone, polyurethane, orpolypropylene, bonded to or molded with the lead 12.

The blades 156 can be deflected toward a distal direction in response toexerting a pulling force on the lead 12 at a threshold axial forcelevel, which is greater than expected day-to-day axial forces. Theblades 156 are sized and configured to yield during proximal passagethrough tissue in result to such forces, causing minimal tissue trauma,and without breaking or leaving fragments, despite the possible presenceof some degree of tissue in-growth. This feature permits the withdrawalof the implanted electrode 16, if desired, by purposeful pulling on thelead 12 at the higher axial force level.

Desirably, the anchoring means 150 is prevented from fully engaging bodytissue until after the electrode 16 has been deployed. The electrode 16is not deployed until after it has been correctly located during theimplantation (installation) process.

More particularly, and as described below, the lead 12 and electrode 16are intended to be percutaneously introduced through a sleeve orintroducer 158 shown in FIG. 36. As shown, the blades 156 assume acollapsed condition against the lead 12 body when within the sleeve 158.In this condition, the blades 156 are shielded from contact with tissue.Once the location is found, the sleeve 158 can be withdrawn, holding thelead 12 and electrode 16 stationary. Free of the sleeve 158, the blades156 spring open to assume their radially deployed condition in tissue,fixing the electrode 16 in the desired location.

The position of the electrode 16 relative to the anchoring means 150,and the use of the sleeve 158, allows for both advancement andretraction of the electrode delivery sleeve 158 during implantationwhile simultaneously delivering test stimulation. During this phase ofthe implantation process, the distal tip of the lead 12 may be exposedto direct tissue contact, or alternatively, the lead 12 may be replacedby a metallic introducing needle that would extend beyond the end of theinsulating delivery sleeve 158. The proximal end of the introducingneedle (or the connector 155 of the lead 12) would be connected to atest stimulator. The sleeve 158 can be drawn back relative to the lead12 to deploy the electrode 16 anchoring means 150, but only when thephysician determines that the desired electrode location has beenreached. The withdrawal of the sleeve 158 from the lead 12 causes theanchoring means 150 to deploy without changing the position of electrode16 in the desired location (or allowing only a small and predictable,set motion of the electrode). Once the sleeve 158 is removed, theflexible, silicone-coated or polyurethane-coat lead 12 and electrode 16are left implanted in the targeted tissue region.

4. Molded Nerve Cuff

In an alternative embodiment, a lead 12 and a cuff electrode 16′ may beused. As FIG. 37 shows, the cuff electrode 16′ includes at least oneelectrically conductive surface 160. It is to be appreciated that thecuff electrode 16′ may be a spiral cuff, as shown, or may also be asplit cylinder cuff. In the illustrated embodiment, there are threeindividually controllable electrically conductive surfaces 160, althoughmore or less may be used. The surface 160 may be solid or the surfacemay be segmented into isolated conductive segments electrically coupledby a wire. It is to be appreciated that additional alternativeconfigurations are possible as well. These surfaces may be manufacturedusing a thin film of metal deposited on a liquid crystal polymersubstrate, or from strips of platinum, for example.

As FIG. 37 shows, the cuff electrode 16′ comprises a body 162 and astrain relief boot 164 that may be molded from a low durometer elastomermaterial (e.g., silicone, such as a two part, translucent, pourablesilicone elastomer, e.g., Nusil MED-4211). The electrically conductivesurfaces 160 are integrated with the body 162 during the moldingprocess. The boot 164 strengthens the junction, to resist the effect oftorque forces that might be applied during implantation and use alongthe lead 12. In addition, the strain relief boot 164 helps to preventtension and/or motion from damaging the lead to cuff interface for alonger flex life.

The molded body 162 of the cuff electrode 16′ is shaped or formed duringthe molding process to normally assume a curled or tubular spiral orrolled configuration. As shown in FIG. 37, in its normal coiledcondition, the body 162 extends in a spiral having a range of greaterthan 360 degrees from end to end, and in one embodiment about 540degrees from end to end. The body 162 can be elastically uncoiled toincrease its inner diameter, i.e., to be initially fitted about theperiphery of a target nerve N, and in response to post-operative changesin the diameter of the target nerve N that might occur due to swelling.The elasticity of the body 162 wraps the electrically conductivesurfaces gently against the periphery of the targeted nerve N. Theelasticity of the body 162 is selected to gently wrap about the targetnerve N without causing damage or trauma. To this end, it is believeddesirable that the elastic memory of the cuff electrode 16′ exhibits apredictable and repeatable pressure vs. diameter relationship thatgradually increases pressure with increase in diameter to allow theelectrode to fit snuggly about the periphery of a nerve, but not tootightly to cause damage (i.e., exerts a maximum pressure about thetarget nerve N that does not exceed about 20 mmHg).

II. Operating System

The implantable pulse generator operating system software 200 (operatingon the microcontroller 36) controls the sequencing and operation of theimplantable pulse generator hardware. As can be seen in FIG. 38, theoperating system software 200 can be broadly grouped into twocategories: the system software 202 and the application software 204.

A. System Software

The system software 202 constitutes a majority of the software codecontrolling the implantable pulse generator 18. As an example, thesystem software may constitute about 85 percent to 95 percent of theoperating system software 200, and, the application software 204 mayconstitute about five percent to fifteen percent of the operating systemsoftware. Structurally, the system software 202 ranges from the lowlevel peripheral drivers 206 that directly interface with theimplantable pulse generator hardware to the higher level softwaredrivers 208 that manages the timing of wireless telemetry communications112 and the encoding and decoding of the wireless messages in accordancewith the communications protocol.

The system software 202 is responsible for monitoring and controllingall the hardware of the implantable pulse generator 18. Key activitiesmay include:

-   -   The activation and disabling of hardware components or        sub-systems as they are required to be functional or are no        longer required. For example, the wireless telemetry hardware is        only enabled when it is required, as a power management        technique. The stimulus power supply is only enabled immediately        before and during the delivery of a stimulus pulse, as a power        management and noise control technique.    -   The generation of precisely timed interrupts or software events.        These software events are used to invoke the application        software 204, update the current time data, and to schedule and        perform regular or periodic “house cleaning” activities and the        interface of system resources, such as, wireless telemetry        communications, time and date information, storage and retrieval        of usage data and operational settings, and monitoring battery        voltage, etc.    -   Configure the wireless telemetry circuitry to “sniff” for any        communications or interference on the wireless telemetry 112.    -   Configure the wireless telemetry circuitry 140 to receive a        command and to send a response.    -   Process any general (not application specific) commands and        generate the associated response (this includes the retrieval of        log data).    -   Generate a stimulus pulse of specified amplitude and pulse        duration.    -   Measure the cathodic phase voltage during a stimulus pulse and        optimize the value of VHH as appropriate.    -   Direct a stimulus pulse to the desired channel(s).    -   Monitor the battery voltage and shut down operations as        necessary in low battery and critical low battery conditions.    -   Monitor the magnitude of the voltage recovered from the power        management (charging) circuitry 130 and the battery voltage to        provide correct information to the implant charger controller        102 (through the wireless telemetry link 112) and to control the        charging process.    -   Measure the value of the VHH power supply and take corrective        actions if necessary.

The system software 204 is also responsible for performing the basicfunctions that are required by all, or most, applications. Thesefunctions may include:

-   -   Invocation of and interface to the application software (code)        204.    -   Making implantable pulse generator and system status information        available to the application software; and similarly, the system        software accepts data generated by the application software and        performs the actions associated with that data (e.g., store        information into non-volatile memory, generate a stimulus pulse        of specified parameters, modify the delay time until the next        stimulus pulse, change status data for subsequent communications        with external hardware, etc.).    -   The execution of the application software on a time or event        scheduled basis (e.g., to be executed every 1/30th second or        whenever a command is received via the wireless telemetry 112).    -   Decode and authenticate (i.e., check for accuracy and        legitimacy) commands received by the wireless telemetry 112.    -   Pass along to the application software any valid, application        specific command received.    -   Encode and transmit any responses made by the application        software    -   Update log entries based on changes to operating modes,        charging, etc.    -   Update log entries in response to data passed by the application        software 204 to the system software 202.

B. Application Software

The application software 204 is implemented as a separate module(s) thatinterfaces with the implantable pulse generator resources (hardware)through calls to software units in the system software 202. This allowsthe application software 204 to be written in relative isolation fromthe details of the implantable pulse generator hardware and the detailsof how the system software 202 manages the hardware. Thus theapplication software 204 utilizes a clearly defined (and limited)interface 203 to the system software 204 and implantable pulse generatorresources (hardware and software) through the use of calls to systemsoftware units (functions).

The application software 204 is responsible for performing theactivities that are specific to the particular application for which theimplantable pulse generator is being used. These functions may include:

-   -   Determining what actions the implantable pulse generator 18 will        take to implement the desired clinical, therapeutic, diagnostic,        or other physiological process for which the implantable pulse        generator was implanted.    -   Defining application status information that will be        communicated to external hardware via the wireless telemetry        112.    -   Determining what usage, history, or diagnostic information        should be stored or retrieved for use by the application or for        telemetry to the external hardware.    -   Establish the stimulus frequency desired. This decision may make        use of the current time information provided by the system        software 202.    -   Establish the amplitude and pulse duration of the next stimulus        pulse to be generated. This decision may also make use of the        current time information provided by the system software.    -   Interpretation of application specific commands received from        the system software 202 and generation of the response to the        application specific commands to the system software.    -   Update entries to any application specific logs.        III. Representative Indications

Due to their technical features, the implantable pulse generator 18 and88 can be used to provide beneficial results in diverse therapeutic andfunctional restorations indications.

For example, in the field of urology, possible indications for use ofthe implantable pulse generators 18 and 68 include the treatment of (i)urinary and fecal incontinence; (ii) micturition/retention; (iii)restoration of sexual function; (iv) defecation/constipation; (v) pelvicfloor muscle activity; and/or (vi) pelvic pain.

The implantable pulse generators 18 and 88 can be used for deep brainstimulation in the treatment of (i) Parkinson's disease; (ii) multiplesclerosis; (iii) essential tremor; (iv) depression; (v) eatingdisorders; (vi) epilepsy; and/or (vii) minimally conscious state.

The implantable pulse generators 18 and 88 can be used for painmanagement by interfering with or blocking pain signals from reachingthe brain, in the treatment of, e.g., (i) peripheral neuropathy; and/or(ii) cancer.

The implantable pulse generators 18 and 88 can be used for vagal nervestimulation for control of epilepsy, depression, or othermood/psychiatric disorders.

The implantable pulse generators 18 and 88 can be used for the treatmentof obstructive sleep apnea.

The implantable pulse generators 18 and 88 can be used for gastricstimulation to prevent reflux or to reduce appetite or food consumption.

The implantable pulse generators 18 and 88 can be used to compensate forvarious cardiac dysfunctions, such as rhythm disorders.

The implantable pulse generators 18 and 88 can be used in functionalrestorations indications such as the restoration of motor control, torestore (i) impaired gait after stroke or spinal cord injury (SCI); (ii)impaired hand and arm function after stroke or SCI; (iii) respiratorydisorders; (iv) swallowing disorders; (v) sleep apnea; and/or (vi)neurotherapeutics, allowing individuals with neurological deficits, suchas stroke survivors or those with multiple sclerosis, to recoverfunctionally.

The foregoing is considered as illustrative only of the principles ofthe invention. Furthermore, since numerous modifications and changeswill readily occur to those skilled in the art, it is not desired tolimit the invention to the exact construction and operation shown anddescribed. While the preferred embodiment has been described, thedetails may be changed without departing from the invention, which isdefined by the claims.

We claim:
 1. A method comprising: transcutaneously recharging arechargeable battery within a housing of an implantable electricalstimulation generator via an externally generated radio frequencymagnetic field, wherein the radio frequency magnetic field is generatedvia a first circuitry of an external controller and a charging coilcoupled to the first circuitry, wherein the externally generated radiofrequency magnetic field is received via a power receiving coil carriedwithin the housing of the implantable electrical stimulation generatorto recharge the rechargeable battery; wirelessly communicating with anantenna and second circuitry of the implantable electrical stimulationgenerator via the first circuitry of the external controller using atleast one of UHF signals or VHF signals, wherein transcutaneouslyrecharging the rechargeable battery comprises transcutaneouslyrecharging the rechargeable battery during a battery recharge period,wherein wirelessly communicating comprises wirelessly receiving, duringthe battery recharge period, status information from the implantableelectrical stimulator to allow the external controller to automaticallyadjust the magnitude of the radio frequency magnetic field; andautomatically adjusting a magnitude of the radio frequency magneticfield based on the received status information.
 2. The method of claim1, further comprising instructing a user to reposition the charging coilbased on the status information.
 3. The method of claim 1, whereinautomatically adjusting the magnitude of the radio frequency magneticfield comprises automatically adjusting the magnitude of the radiofrequency magnetic field up to about 300 percent of an initialmagnitude.
 4. The method of claim 1, wherein the status informationincludes an indication of a battery charge status and an indication of amagnitude of power recovered by the power receiving coil.
 5. The methodof claim 1, wherein the status information includes a magnitude of avoltage of the battery.
 6. The method of claim 1, wherein the radiofrequency magnetic field comprises a frequency between about 30 KHz andabout 300 KHz.
 7. The method of claim 1, wherein wirelesslycommunicating with the antenna and the first circuitry of theimplantable electrical stimulation generator via the second circuitry ofthe external controller using at least one of the UHF signals or the VHFsignals comprises wirelessly communicating with the antenna and thefirst circuitry of the implantable electrical stimulation generator viathe second circuitry of the external controller using signals in theMedical Implant Communications Service (MICS) band between about 402 MHzand about 405 MHz.
 8. The method of claim 1, wherein wirelesslycommunicating comprises wirelessly receiving, from the implantableelectrical stimulation generator, instructions to increase or decrease astrength of the magnetic field during the recharging of the battery,while at the same time, generating the radio frequency magnetic field torecharge the rechargeable battery.
 9. The method of claim 1, wherein theantenna is carried within the housing of the implantable electricalstimulation generator.
 10. The method of claim 1, wherein an outsidediameter of the charging coil is between about 40 millimeters and about70 millimeters, and a thickness of the charging coil as measuredperpendicular to the diameter of the charging coil is between aboutthree millimeters and about eleven millimeters.
 11. The method of claim1, wherein automatically adjusting the magnitude of the radio frequencymagnetic field based on the received status information comprisesautomatically increasing the magnitude of the radio frequency magneticfield based on the received status information.